Microfluidic free interface diffusion techniques

ABSTRACT

A static fluid and a second fluid are placed into contact along a microfluidic free interface and allowed to mix by diffusion without convective flow across the interface. In accordance with one embodiment of the present invention, the fluids are static and initially positioned on either side of a closed valve structure in a microfluidic channel having a width that is tightly constrained in at least one dimension. The valve is then opened, and no-slip layers at the sides of the microfluidic channel suppress convective mixing between the two fluids along the resulting interface. Applications for microfluidic free interfaces in accordance with embodiments of the present invention include, but are not limited to, protein crystallization studies, protein solubility studies, determination of properties of fluidics systems, and a variety of biological assays such as diffusive immunoassays, substrate turnover assays, and competitive binding assays.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No. 12/762,170, filed Apr. 16, 2010, which is a divisional of U.S. patent application Ser. No. 12/001,768, filed Dec. 11, 2007, issued as U.S. Pat. No. 7,704,322 on Apr. 27, 2010, which is a divisional of U.S. patent application Ser. No. 10/265,473, filed Oct. 4, 2002, issued as U.S. Pat. No. 7,306,672 on Dec. 11, 2007, which is a continuation-in-part of U.S. patent application Ser. No. 09/826,583, filed Apr. 6, 2001, issued as U.S. Pat. No. 6,899,137 on May 31, 2005, and is a continuation-in-part of U.S. patent application Ser. No. 09/887,997, filed Jun. 22, 2001, issued as U.S. Pat. No. 7,052,545 on May 30, 2006, and is a continuation-in-part of U.S. patent application Ser. No. 10/117,978, filed Apr. 5, 2002, now issued as U.S. Pat. No. 7,195,670 on Mar. 27, 2007. These prior patent applications are hereby incorporated by reference for all purposes.

STATEMENT AS TO RIGHTS TO INVENTIONS MADE UNDER FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made with government support under Grant No. HG-01642-02 awarded by the National Institutes of Health. The government has certain rights in the invention.

BACKGROUND OF THE INVENTION

Free-interface diffusion techniques have been employed in a number of scientific applications, most notably in protein crystallization studies. In a free interface, mixing between different fluids occurs entirely as a result of diffusion. The creation of such free interfaces offer the advantage of gradual and non-directional mixing, avoiding asymmetries and steep concentration gradients associated with convective flow between the fluids.

However, certain technical difficulties have rendered conventional macroscopic free interfaces unsuitable for high throughput screening applications, and they are not currently widely used in the crystallographic community for several reasons.

First, conventional macroscopic fluidic interfaces are established by dispensing solutions into a narrow container; such as a capillary tube having a width of 100 μm or greater, or a deep well in a culture plate. FIGS. 30A-B show simplified cross-sectional views of the attempted formation of a macroscopic free-interface in a capillary tube 9300. The act of dispensing a second fluid 9302 into a first fluid 9304 may give rise to convective mixing and result in a poorly defined fluidic interface 9306.

Unfortunately, the fluids may not be sucked into a capillary serially to eliminate this problem. FIGS. 31A-B show the mixing, between a first solution 9400 and a second solution 9402 in a capillary tube 9404 that would result due to the parabolic velocity distribution of pressure driven convective Poiseuille flow, also resulting in a poorly defined fluidic interface 9406.

Moreover, the container for a macroscopic free-interface crystallization regime must have dimensions making them accessible to a pipette tip, needle, capillary or other g tool, and necessitating the use of relatively large (10-100 μl) fluid volumes.

In order to avoid unwanted convective mixing during macroscopic free interface diffusion experiments, considerable care must be exercised both during dispensing of the fluids and during the diffusion period. For this reason cumbersome protocols are often used to define a macroscopic free-interface. In one conventional approach, one fluid may be frozen or otherwise converted to solid phase prior to addition of the second fluid. In an alternative conventional approach, the one fluid containing the sample is converted into a solid through polymerization, trapping the sample within the polymer, which is then exposed to the second fluid. See generally Garcia-Ruiz et al., “Agarose as Crystallization Media for Proteins I: Transport Processes”, J. Crystal Growth 232, 165-172 (2001).

The problems of convective mixing associated with macroscopic free interface diffusion studies is compounded in that two fluids of differing density will mix by gravity induced convection if they are not stored at the proper orientation, additionally complicating the storage of reactions. This is shown in FIGS. 32A-C, wherein over time first solution 9500 having a density greater than the density of second solution 9502 merely sinks to form a static bottom layer 9504 that is not conducive to formation of a diffusion gradient along the length of a capillary tube.

Other approaches to the formation of free interfaces have focused upon microfluidic structures. For example, Kamholz et al., “Quantitative Analysis of Molecular Interaction in a Microfluidic Channel: the T-Sensor”, Anal. Chem. Vol. 71, No. 23 (1999), describe the formation of a diffusive interface between two flowing fluids inlet on sides of microfluidic T-shaped junction and then moving together along the stem. The dimensions of the microfluidic channels impose laminar flow on the fluids, such that convective mixing between them is eliminated and diffusion only occurs across the interface.

While the diffusion described by Kamholz et al. may be useful, it offers the distinct disadvantage of necessarily creating an elongated interface between the fluids, such that substantial volumes of the fluids are continuously exposed to the steep concentration gradient occurring at the interface between them. Such a steep gradient poses a number of disadvantages. In the context of a protein crystallization experimental protocol, one such disadvantage is the possibility of solvent shock and the precipitation of sample in non-crystalline form along the interface.

Accordingly, there is a need in the art for methods and structures for accomplishing diffusive introduction of material under highly controlled conditions that limit the exposure of volumes of solution to steep concentration gradients.

BRIEF SUMMARY OF THE INVENTION

A static fluid and a second fluid are placed into contact along a microfluidic free interface and allowed to mix by diffusion without convective flow across the interface. In accordance with one embodiment of the present invention, the fluids are static and initially positioned on either side of a closed valve structure in a microfluidic channel having a width that is tightly constrained in at least one dimension. The valve is then opened, and no-slip layers at the sides of the microfluidic channel suppress convective mixing between the two fluids along the resulting interface. Applications for microfluidic free interfaces in accordance with embodiments of the present invention include, but are not limited to, protein crystallization studies, protein solubility studies, determination of properties of fluidics systems, and a variety of biological assays such as diffusive immunoassays, substrate turnover assays, and competitive binding assays.

An embodiment of a method of mixing two fluids in accordance with the present invention comprises disposing a first static fluid, disposing a second fluid proximate to the first static fluid to form a fluidic interface, and suppressing convective flow of the first and second fluids such that mixing between the first and second fluids across the interface occurs substantially exclusively by diffusion.

An alternative embodiment of a method of mixing two fluids in accordance with the present invention comprises providing a first static fluid having a total volume, providing a second static fluid having a total volume, disposing a portion of the first static fluid in contact with a portion of the second static fluid to form a fluidic interface, such that a minimum volume of the first and second fluids is exposed to a steepest concentration gradient present immediately along the fluidic interface.

An embodiment of a method of determining a property of a fluidic system in accordance with the present invention comprises disposing a first static fluid containing a target, disposing a second static fluid proximate to the first fluid to form a fluidic interface, and suppressing convective flow of the first and second fluids such that mixing occurs across the interface solely by diffusion to reveal the physical property.

An embodiment of a structure for analyzing a property of a target material comprises a first microfabricated region configured to contain a volume of a first static fluid including a target, a second microfabricated region configured to contain a volume of a second static fluid including an analyte known to bind to the target, and a valve actuable to place the volume of the first fluid in diffusive communication with the volume of the second fluid across a free microfluidic interface. A detector is configured to detect a presence of the analyte in the first microfabricated region.

An embodiment of a method of reducing a concentration of a small molecule in a protein sample in accordance with the present invention comprises disposing a first static fluid containing the protein sample in a microfluidic channel, disposing a second static fluid having a low concentration of the small molecule proximate to the first fluid to form a fluidic interface, and suppressing convective flow of the first and second fluids such that the small molecule from the first static fluid diffuses across the interface solely by diffusion to reduce the concentration of small molecule present in the first static fluid.

An embodiment of a method for determining reaction between a ligand and a target comprises positioning a first fluid containing the ligand in a first chamber, positioning a second fluid containing the target in a second chamber, and establishing a microfluidic free interface between the first and second fluids in a channel connecting the first and the second chamber. Mixing is allowed to occur by diffusion across the microfluidic free interface between the first and second fluids, such that the ligand binds to the target and reactivity between the ligand and target can be determined by deviation of at least one of a temporal diffusion profile and a spatial diffusion profile from a corresponding profile expected in the absence of the target.

An embodiment of a method of sampling a chemical reaction under a spectrum of conditions in accordance with the present invention comprises forming a microfluidic free interface between a first fluid containing a first reactant and a second fluid containing a second reactant, and causing diffusion of the first reactant into the second fluid to create a concentration gradient of the first reactant. Reaction between the first reactant at various concentrations along the gradient and the second reactant is observed.

An embodiment of a method of creating a concentration gradient of a chemical species in accordance with the present invention comprises disposing a first fluid containing the chemical species, and disposing a second static fluid proximate to the first static fluid to form a microfluidic free interface. Convective flow of the first and second fluids is suppressed such that mixing between the first and second fluids across the microfluidic free interface occurs substantially exclusively by diffusion and a concentration gradient of the chemical species is created.

An embodiment of a method of introducing a drug into a subject in accordance with the present invention comprises disposing a first static fluid containing the drug in a microfluidic device, and disposing a second fluid in the microfluidic device between the first static fluid and the subject. A microfluidic free interface is established between the first fluid and the second fluid, such that a predetermined amount of the drug diffuses to reach the subject only after a predetermined time.

An embodiment of a method of controlling the concentration of reactants during a chemical reaction in accordance with the present invention comprises disposing a first static fluid containing a first reactant in a microfluidic structure, and disposing a second fluid in the microfluidic structure proximate to the first static fluid, the second static fluid containing a second reactant. A microfluidic free interface is established between the first and second fluids, and diffusion between the first and second fluids across the microfluidic free interface is allowed to determine the relative concentration of the first and second reactants.

These and other embodiments of the present invention, as well as its advantages and features, are described in more detail in conjunction with the text below and attached figures.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an illustration of a first elastomeric layer formed on top of a micromachined mold.

FIG. 2 is an illustration of a second elastomeric layer formed on top of a micromachined mold.

FIG. 3 is an illustration of the elastomeric layer of FIG. 2 removed from the micromachined mold and positioned over the top of the elastomeric layer of FIG. 1.

FIG. 4 is an illustration corresponding to FIG. 3, but showing the second elastomeric layer positioned on top of the first elastomeric layer.

FIG. 5 is an illustration corresponding to FIG. 4, but showing the first and second elastomeric layers bonded together.

FIG. 6 is an illustration corresponding to FIG. 5, but showing the first micromachined mold removed and a planar substrate positioned in its place.

FIG. 7A is an illustration corresponding to FIG. 6, but showing the elastomeric structure sealed onto the planar substrate.

FIG. 7B is a front sectional view corresponding to FIG. 7A, showing an open flow channel.

FIGS. 7C-7G are illustrations showing steps of a method for forming an elastomeric structure having a membrane formed from a separate elastomeric layer.

FIG. 7H is a front sectional view showing the valve of FIG. 7B in an actuated state.

FIGS. 8A and 8B illustrates valve opening vs. applied pressure for various flow channels.

FIG. 9 illustrates time response of a 100 μm×100 μm×10 μm RTV microvalve.

FIG. 10 is a front sectional view of the valve of FIG. 7B showing actuation of the membrane.

FIG. 11 is a front sectional view of an alternative embodiment of a valve having a flow channel with a curved upper surface.

FIG. 12A is a top schematic view of an on/off valve.

FIG. 12B is a sectional elevation view along line 23B-23B in FIG. 12A

FIG. 13A is a top schematic view of a peristaltic pumping system.

FIG. 13B is a sectional elevation view along line 24B-24B in FIG. 13A

FIG. 14 is a graph showing experimentally achieved pumping rates vs. frequency for an embodiment of the peristaltic pumping system of FIG. 13.

FIG. 15A is a top schematic view of one control line actuating multiple flow lines simultaneously.

FIG. 15B is a sectional elevation view along line 26B-26B in FIG. 15A.

FIG. 16 is a schematic illustration of a multiplexed system adapted to permit flow through various channels.

FIG. 17A is a plan view of a flow layer of an addressable reaction chamber structure.

FIG. 17B is a bottom plan view of a control channel layer of an addressable reaction chamber structure.

FIG. 17C is an exploded perspective view of the addressable reaction chamber structure formed by bonding the control channel layer of FIG. 17B to the top of the flow layer of FIG. 17A.

FIG. 17D is a sectional elevation view corresponding to FIG. 17C, taken along line 28D-28D in FIG. 17C.

FIG. 18 is a schematic of a system adapted to selectively direct fluid flow into any of an array of reaction wells.

FIG. 19 is a schematic of a system adapted for selectable lateral flow between parallel flow channels.

FIG. 20A is a bottom plan view of first layer (i.e.: the flow channel layer) of elastomer of a switchable flow array.

FIG. 20B is a bottom plan view of a control channel layer of a switchable flow array.

FIG. 20C shows the alignment of the first layer of elastomer of FIG. 20A with one set of control channels in the second layer of elastomer of FIG. 20B.

FIG. 20D also shows the alignment of the first layer of elastomer of FIG. 20A with the other set of control channels in the second layer of elastomer of FIG. 20B.

FIGS. 21A-J show views of one embodiment of a normally-closed valve structure in accordance with the present invention.

FIGS. 22A and 22B show plan views illustrating operation of one embodiment of a side-actuated valve structure in accordance with the present invention.

FIG. 23 shows a cross-sectional view of one embodiment of a composite structure in accordance with the present invention.

FIG. 24 shows a cross-sectional view of another embodiment of a composite structure in accordance with the present invention.

FIG. 25 shows a cross-sectional view of another embodiment of a composite structure in accordance with the present invention.

FIGS. 26A-D show plan views illustrating operation of one embodiment of a cell pen structure in accordance with the present invention.

FIGS. 27A-B show plan and cross-sectional views illustrating operation of one embodiment of a cell cage structure in accordance with the present invention.

FIGS. 28A-B show plan views of operation of a wiring structure utilizing cross-channel injection in accordance with the embodiment of the present invention.

FIGS. 29A-D illustrate cross-sectional views of metering by volume exclusion in accordance with an embodiment of the present invention.

FIGS. 30A-B show simplified cross-sectional views of the attempted formation of a macroscopic free-interface in a capillary tube.

FIGS. 31A-B show simplified cross-sectional views of convective mixing between a first solution and a second solution in a capillary tube resulting from a parabolic velocity distribution of pressure driven Poiseuille flow.

FIGS. 32A-C show simplified cross-sectional views of interaction in a capillary tube between a first solution having a density greater than the density of second solution.

FIG. 33A shows a simplified cross-sectional view of a microfluidic free interface in accordance with an embodiment of the present invention.

FIG. 33B shows a simplified cross-sectional view of a conventional non-microfluidic interface.

FIGS. 34A-D show plan views of the priming of a flow channel and formation of a microfluidic free interface in accordance with an embodiment of the present invention.

FIGS. 35A-E show simplified schematic views of the use of “break-through” valves to create a microfluidic free interface.

FIGS. 36A-D are simplified schematic diagrams plotting concentration versus distance for two fluids in diffusing across a microfluidic free interface in accordance with an embodiment of the present invention.

FIG. 37A shows a simplified plan view of a flow channel overlapped at intervals by a forked control channel to define a plurality of chambers (A-G) positioned on either side of a separately-actuated interface valve.

FIGS. 37B-D plot solvent concentration at different times for the flow channel shown in FIG. 37A.

FIG. 38A shows three sets of pairs of chambers connected by microchannels of a different length.

FIG. 38B plots equilibration time versus channel length.

FIG. 39 shows four pairs of chambers, each having different arrangements of connecting microchannel(s).

FIG. 40A shows a plan view of a simple embodiment of a microfluidic structure in accordance with the present invention.

FIG. 40B is a simplified plot of concentration versus distance for the structure of FIG. 40A.

FIG. 41 plots the time required for the concentration in one of the reservoirs of FIG. 40A to reach 0.6 of the final equilibration concentration, versus channel length.

FIG. 42 plots the inverse of the time required for the concentration in one of the reservoirs to reach 0.6 of the final equilibration concentration (T_(0.6)), versus the area of the fluidic interface of FIG. 40A.

FIG. 43 presents a phase diagram depicting the phase space between fluids A and B, and the path in phase space traversed in the reservoirs as the fluids diffuse across the microfluidic free interface of FIG. 40A.

FIG. 44A shows a simplified schematic view of a protein crystal being formed utilizing a conventional macroscopic free interface diffusion technique.

FIG. 44B shows a simplified schematic view of a protein crystal being formed utilizing diffusion across a microfluidic free interface in accordance with an embodiment of the present invention.

FIG. 45A shows a simplified plan view of one embodiment of a microfluidic structure for performing diffusion analysis in accordance with the present invention.

FIG. 45B plots expected and observed signal intensity versus diffusion time for the structure shown in FIG. 45A.

FIG. 46A shows a simplified plan view of an alternative embodiment of a microfluidic structure for performing diffusion analysis in accordance with the present invention.

FIG. 46B plots expected and observed signal intensity versus diffusion distance for the structure shown in FIG. 46A.

FIG. 47 shows a plan view of the creation of a microfluidic free interface between a flowing fluid and a dead-ended branch channel.

FIG. 48 shows a plan view of diffusion of nutrients and waste across a microfluidic free interface to and from a cell cage.

FIG. 49 shows a simplified plan view of one embodiment of a microfluidic structure in accordance with the present invention for performing a competitive binding assay.

FIG. 50 shows a simplified plan view of an alternative embodiment of a microfluidic structure in accordance with the present invention for performing a competitive binding assay.

FIG. 51 shows a simplified plan view of an embodiment of a microfluidic structure in accordance with the present invention for performing a substrate turnover assay.

FIG. 52 shows a simplified plan view of one embodiment of a microfluidic structure in accordance with the present invention for creating diffusion gradients of two different species in different dimensions.

FIG. 53 shows a simplified plan view of an alternative embodiment of a microfluidic structure in accordance with the present invention for creating diffusion gradients of two different species in different dimensions.

FIG. 54 plots Log(R/B) vs. number of slugs injected for one embodiment of a cross-flow injection system in accordance with the present invention.

DETAILED DESCRIPTION OF THE INVENTION

Portions of the prior discussion have referenced the use of microfluidic structures to accomplish free interface diffusion in accordance with embodiments of the present invention. The following discussion relates to formation of microfabricated fluidic devices utilizing elastomer materials, as described generally in U.S. patent application Ser. No. 09/826,585 filed Apr. 6, 2001, Ser. No. 09/724,784 filed Nov. 28, 2000, and Ser. No. 09/605,520, filed Jun. 27, 2000. These patent applications are hereby incorporated by reference.

I. Microfabricated Elastomer Structures

1. Methods of Fabricating

Exemplary methods of fabricating the present invention are provided herein. It is to be understood that the present invention is not limited to fabrication by one or the other of these methods. Rather, other suitable methods of fabricating the present microstructures, including modifying the present methods, are also contemplated.

FIGS. 1 to 7B illustrate sequential steps of a first preferred method of fabricating the present microstructure, (which may be used as a pump or valve). FIGS. 8 to 18 illustrate sequential steps of a second preferred method of fabricating the present microstructure, (which also may be used as a pump or valve).

As will be explained, the preferred method of FIGS. 1 to 7B involves using pre-cured elastomer layers which are assembled and bonded. In an alternative method, each layer of elastomer may be cured “in place”. In the following description “channel” refers to a recess in the elastomeric structure which can contain a flow of fluid or gas.

Referring to FIG. 1, a first micro-machined mold 10 is provided. Micro-machined mold 10 may be fabricated by a number of conventional silicon processing methods, including but not limited to photolithography, ion-milling, and electron beam lithography.

As can be seen, micro-machined mold 10 has a raised line or protrusion 11 extending there along. A first elastomeric layer 20 is cast on top of mold 10 such that a first recess 21 will be formed in the bottom surface of elastomeric layer 20, (recess 21 corresponding in dimension to protrusion 11), as shown.

As can be seen in FIG. 2, a second micro-machined mold 12 having a raised protrusion 13 extending there along is also provided. A second elastomeric layer 22 is cast on top of mold 12, as shown, such that a recess 23 will be formed in its bottom surface corresponding to the dimensions of protrusion 13.

As can be seen in the sequential steps illustrated in FIGS. 3 and 4, second elastomeric layer 22 is then removed from mold 12 and placed on top of first elastomeric layer 20. As can be seen, recess 23 extending along the bottom surface of second elastomeric layer 22 will form a flow channel 32.

Referring to FIG. 5, the separate first and second elastomeric layers 20 and 22 (FIG. 4) are then bonded together to form an integrated (i.e.: monolithic) elastomeric structure 24.

As can been seen in the sequential step of FIGS. 6 and 7A, elastomeric structure 24 is then removed from mold 10 and positioned on top of a planar substrate 14. As can be seen in FIGS. 7A and 7B, when elastomeric structure 24 has been sealed at its bottom surface to planar substrate 14, recess 21 will form a flow channel 30.

The present elastomeric structures form a reversible hermetic seal with nearly any smooth planar substrate. An advantage to forming a seal this way is that the elastomeric structures may be peeled up, washed, and re-used. In preferred aspects, planar substrate 14 is glass. A further advantage of using glass is that glass is transparent, allowing optical interrogation of elastomer channels and reservoirs. Alternatively, the elastomeric structure may be bonded onto a flat elastomer layer by the same method as described above, forming a permanent and high-strength bond. This may prove advantageous when higher back pressures are used.

As can be seen in FIGS. 7A and 7B, flow channels 30 and 32 are preferably disposed at an angle to one another with a small membrane 25 of substrate 24 separating the top of flow channel 30 from the bottom of flow channel 32.

In preferred aspects, planar substrate 14 is glass. An advantage of using glass is that the present elastomeric structures may be peeled up, washed and reused. A further advantage of using glass is that optical sensing may be employed. Alternatively, planar substrate 14 may be an elastomer itself, which may prove advantageous when higher back pressures are used.

The method of fabrication just described may be varied to form a structure having a membrane composed of an elastomeric material different than that forming the walls of the channels of the device. This variant fabrication method is illustrated in FIGS. 7C-G.

Referring to FIG. 7C, a first micro-machined mold 10 is provided. Micro-machined mold 10 has a raised line or protrusion 11 extending there along. In FIG. 7D, first elastomeric layer 20 is cast on top of first micro-machined mold 10 such that the top of the first elastomeric layer 20 is flush with the top of raised line or protrusion 11. This may be accomplished by carefully controlling the volume of elastomeric material spun onto mold 10 relative to the known height of raised line 11. Alternatively, the desired shape could be formed by injection molding.

In FIG. 7E, second micro-machined mold 12 having a raised protrusion 13 extending there along is also provided. Second elastomeric layer 22 is cast on top of second mold 12 as shown, such that recess 23 is formed in its bottom surface corresponding to the dimensions of protrusion 13.

In FIG. 7F, second elastomeric layer 22 is removed from mold 12 and placed on top of third elastomeric layer 222. Second elastomeric layer 22 is bonded to third elastomeric layer 20 to form integral elastomeric block 224 using techniques described in detail below. At this point in the process, recess 23 formerly occupied by raised line 13 will form flow channel 23.

In FIG. 7G, elastomeric block 224 is placed on top of first micro-machined mold 10 and first elastomeric layer 20. Elastomeric block and first elastomeric layer 20 are then bonded together to form an integrated (i.e.: monolithic) elastomeric structure 24 having a membrane composed of a separate elastomeric layer 222.

When elastomeric structure 24 has been sealed at its bottom surface to a planar substrate in the manner described above in connection with FIG. 7A, the recess formerly occupied by raised line 11 will form flow channel 30.

The variant fabrication method illustrated above in conjunction with FIGS. 7C-G offers the advantage of permitting the membrane portion to be composed of a separate material than the elastomeric material of the remainder of the structure. This is important because the thickness and elastic properties of the membrane play a key role in operation of the device. Moreover, this method allows the separate elastomer layer to readily be subjected to conditioning prior to incorporation into the elastomer structure. As discussed in detail below, examples of potentially desirable condition include the introduction of magnetic or electrically conducting species to permit actuation of the membrane, and/or the introduction of dopant into the membrane in order to alter its elasticity.

While the above method is illustrated in connection with forming various shaped elastomeric layers formed by replication molding on top of a micromachined mold, the present invention is not limited to this technique. Other techniques could be employed to form the individual layers of shaped elastomeric material that are to be bonded together. For example, a shaped layer of elastomeric material could be formed by laser cutting or injection molding, or by methods utilizing chemical etching and/or sacrificial materials as discussed below in conjunction with the second exemplary method.

An alternative method fabricates a patterned elastomer structure utilizing development of photoresist encapsulated within elastomer material. However, the methods in accordance with the present invention are not limited to utilizing photoresist. Other materials such as metals could also serve as sacrificial materials to be removed selective to the surrounding elastomer material, and the method would remain within the scope of the present invention. For example, gold metal may be etched selective to RTV 615 elastomer utilizing the appropriate chemical mixture.

2. Layer and Channel Dimensions

Microfabricated refers to the size of features of an elastomeric structure fabricated in accordance with an embodiment of the present invention. In general, variation in at least one dimension of microfabricated structures is controlled to the micron level, with at least one dimension being microscopic (i.e. below 1000 μm). Microfabrication typically involves semiconductor or MEMS fabrication techniques such as photolithography and spincoating that are designed for to produce feature dimensions on the microscopic level, with at least some of the dimension of the microfabricated structure requiring a microscope to reasonably resolve/image the structure.

In preferred aspects, flow channels 30, 32, 60 and 62 preferably have width-to-depth ratios of about 10:1. A non-exclusive list of other ranges of width-to-depth ratios in accordance with embodiments of the present invention is 0.1:1 to 100:1, more preferably 1:1 to 50:1, more preferably 2:1 to 20:1, and most preferably 3:1 to 15:1. In an exemplary aspect, flow channels 30, 32, 60 and 62 have widths of about 1 to 1000 microns. A non-exclusive list of other ranges of widths of flow channels in accordance with embodiments of the present invention is 0.01 to 1000 microns, more preferably 0.05 to 1000 microns, more preferably 0.2 to 500 microns, more preferably 1 to 250 microns, and most preferably 10 to 200 microns. Exemplary channel widths include 0.1 μm, 1 μm, 2 μm, 5 μm, 10 μm, 20 μm, 30 μm, 40 μm, 50 μm, 60 μm, 70 μm, 80 μm, 90 μm, 100 μm, 110 μm, 120 μm, 130 μm, 140 μm, 150 μm, 160 μm, 170 μm, 180 μm, 190 μm, 200 μm, 210 μm, 220 μm, 230 μm, 240 μm, and 250 μm.

Flow channels 30, 32, 60, and 62 have depths of about 1 to 100 microns. A non-exclusive list of other ranges of depths of flow channels in accordance with embodiments of the present invention is 0.01 to 1000 microns, more preferably 0.05 to 500 microns, more preferably 0.2 to 250 microns, and more preferably 1 to 100 microns, more preferably 2 to 20 microns, and most preferably 5 to 10 microns. Exemplary channel depths include including 0.01 μm, 0.02 μm, 0.05 μm, 0.11 μm, 0.2 μm, 0.5 μm, 11 μm, 2 μm, 3 μm, 4 μm, 5 μm, 7.5 μm, 10 μm, 12.5 μm, 15 μm, 17.5 μm, 20 μm, 22.5 μm, 25 μm, 30 μm, 40 μm, 50 μm, 75 μm, 100 μm, 150 μm, 200 μm, and 250 μm.

The flow channels are not limited to these specific dimension ranges and examples given above, and may vary in width in order to affect the magnitude of force required to deflect the membrane as discussed at length below in conjunction with FIG. 27. For example, extremely narrow flow channels having a width on the order of 0.01 μm may be useful in optical and other applications, as discussed in detail below. Elastomeric structures which include portions having channels of even greater width than described above are also contemplated by the present invention, and examples of applications of utilizing such wider flow channels include fluid reservoir and mixing channel structures.

The Elastomeric layers may be cast thick for mechanical stability. In an exemplary embodiment, elastomeric layer 22 of FIG. 1 is 50 microns to several centimeters thick, and more preferably approximately 4 mm thick. A non-exclusive list of ranges of thickness of the elastomer layer in accordance with other embodiments of the present invention is between about 0.1 micron to 10 cm, 1 micron to 5 cm, 10 microns to 2 cm, 100 microns to 10 mm.

Accordingly, membrane 25 of FIG. 7B separating flow channels 30 and 32 has a typical thickness of between about 0.01 and 1000 microns, more preferably 0.05 to 500 microns, more preferably 0.2 to 250, more preferably 1 to 100 microns, more preferably 2 to 50 microns, and most preferably 5 to 40 microns. As such, the thickness of elastomeric layer 22 is about 100 times the thickness of elastomeric layer 20. Exemplary membrane thicknesses include 0.01 μm, 0.02 μm, 0.03 μm, 0.05 μm, 0.1 μm, 0.2 μm, 0.3 μm, 0.5 μm, 1 μm, 2 μm, 3 μm, 5 μm, 7.5 μm, 10 μm, 12.5 μm, 15 μm, 17.5 μm, 20 μm, 22.5 μm, 25 μm, 30 μm, 40 μm, 50 μm, 75 μm, 100 μm, 150 μm, 200 μm, 250 μm, 300 μm, 400 μm, 500 μm, 750 μm, and 1000 μm.

3. Soft Lithographic Bonding

Preferably, elastomeric layers are bonded together chemically, using chemistry that is intrinsic to the polymers comprising the patterned elastomer layers. Most preferably, the bonding comprises two component “addition cure” bonding.

In a preferred aspect, the various layers of elastomer are bound together in a heterogeneous bonding in which the layers have a different chemistry. Alternatively, a homogenous bonding may be used in which all layers would be of the same chemistry. Thirdly, the respective elastomer layers may optionally be glued together by an adhesive instead. In a fourth aspect, the elastomeric layers may be thermoset elastomers bonded together by heating.

In one aspect of homogeneous bonding, the elastomeric layers are composed of the same elastomer material, with the same chemical entity in one layer reacting with the same chemical entity in the other layer to bond the layers together. In one embodiment, bonding between polymer chains of like elastomer layers may result from activation of a crosslinking agent due to light, heat, or chemical reaction with a separate chemical species.

Alternatively in a heterogeneous aspect, the elastomeric layers are composed of different elastomeric materials, with a first chemical entity in one layer reacting with a second chemical entity in another layer. In one exemplary heterogeneous aspect, the bonding process used to bind respective elastomeric layers together may comprise bonding together two layers of RTV 615 silicone. RTV 615 silicone is a two-part addition-cure silicone rubber. Part A contains vinyl groups and catalyst; part B contains silicon hydride (Si—H) groups. The conventional ratio for RTV 615 is 10A:1B. For bonding, one layer may be made with 30A:1B (i.e. excess vinyl groups) and the other with 3A:1B (i.e. excess Si—H groups). Each layer is cured separately. When the two layers are brought into contact and heated at elevated temperature, they bond irreversibly forming a monolithic elastomeric substrate.

In an exemplary aspect of the present invention, elastomeric structures are formed utilizing Sylgard 182, 184 or 186, or aliphatic urethane diacrylates such as (but not limited to) Ebecryl 270 or Irr 245 from UCB Chemical.

In one embodiment in accordance with the present invention, two-layer elastomeric structures were fabricated from pure acrylated Urethane Ebe 270. A thin bottom layer was spin coated at 8000 rpm for 15 seconds at 170° C. The top and bottom layers were initially cured under ultraviolet light for 10 minutes under nitrogen utilizing a Model ELC 500 device manufactured by Electrolite corporation. The assembled layers were then cured for an additional 30 minutes. Reaction was catalyzed by a 0.5% vol/vol mixture of Irgacure 500 manufactured by Ciba-Geigy Chemicals. The resulting elastomeric material exhibited moderate elasticity and adhesion to glass.

In another embodiment in accordance with the present invention, two-layer elastomeric structures were fabricated from a combination of 25% Ebe 270/50% Irr245/25% isopropyl alcohol for a thin bottom layer, and pure acrylated Urethane Ebe 270 as a top layer. The thin bottom layer was initially cured for 5 min, and the top layer initially cured for 10 minutes, under ultraviolet light under nitrogen utilizing a Model ELC 500 device manufactured by Electrolite corporation. The assembled layers were then cured for an additional 30 minutes. Reaction was catalyzed by a 0.5% vol/vol mixture of Irgacure 500 manufactured by Ciba-Geigy Chemicals. The resulting elastomeric material exhibited moderate elasticity and adhered to glass.

Alternatively, other bonding methods may be used, including activating the elastomer surface, for example by plasma exposure, so that the elastomer layers/substrate will bond when placed in contact. For example, one possible approach to bonding together elastomer layers composed of the same material is set forth by Duffy et al, “Rapid Prototyping of Microfluidic Systems in Poly(dimethylsiloxane)”, Analytical Chemistry (1998), 70, 4974-4984, incorporated herein by reference. This paper discusses that exposing polydimethylsiloxane (PDMS) layers to oxygen plasma causes oxidation of the surface, with irreversible bonding occurring when the two oxidized layers are placed into contact.

Yet another approach to bonding together successive layers of elastomer is to utilize the adhesive properties of uncured elastomer. Specifically, a thin layer of uncured elastomer such as RTV 615 is applied on top of a first cured elastomeric layer. Next, a second cured elastomeric layer is placed on top of the uncured elastomeric layer. The thin middle layer of uncured elastomer is then cured to produce a monolithic elastomeric structure. Alternatively, uncured elastomer can be applied to the bottom of a first cured elastomer layer, with the first cured elastomer layer placed on top of a second cured elastomer layer. Curing the middle thin elastomer layer again results in formation of a monolithic elastomeric structure.

Where encapsulation of sacrificial layers is employed to fabricate the elastomer structure, bonding of successive elastomeric layers may be accomplished by pouring uncured elastomer over a previously cured elastomeric layer and any sacrificial material patterned thereupon. Bonding between elastomer layers occurs due to interpenetration and reaction of the polymer chains of an uncured elastomer layer with the polymer chains of a cured elastomer layer. Subsequent curing of the elastomeric layer will create a bond between the elastomeric layers and create a monolithic elastomeric structure.

Referring to the first method of FIGS. 1 to 7B, first elastomeric layer 20 may be created by spin-coating an RTV mixture on microfabricated mold 12 at 2000 rpm's for 30 seconds yielding a thickness of approximately 40 microns. Second elastomeric layer 22 may be created by spin-coating an RTV mixture on microfabricated mold 11. Both layers 20 and 22 may be separately baked or cured at about 80° C. for 1.5 hours. The second elastomeric layer 22 may be bonded onto first elastomeric layer 20 at about 80° C. for about 1.5 hours.

Micromachined molds 10 and 12 may be patterned photoresist on silicon wafers. In an exemplary aspect, a Shipley SJR 5740 photoresist was spun at 2000 rpm patterned with a high resolution transparency film as a mask and then developed yielding an inverse channel of approximately 10 microns in height. When baked at approximately 200° C. for about 30 minutes, the photoresist reflows and the inverse channels become rounded. In preferred aspects, the molds may be treated with trimethylchlorosilane (TMCS) vapor for about a minute before each use in order to prevent adhesion of silicone rubber.

4. Suitable Elastomeric Materials

Allcock et al, Contemporary Polymer Chemistry, 2^(nd) Ed. describes elastomers in general as polymers existing at a temperature between their glass transition temperature and liquefaction temperature. Elastomeric materials exhibit elastic properties because the polymer chains readily undergo torsional motion to permit uncoiling of the backbone chains in response to a force, with the backbone chains recoiling to assume the prior shape in the absence of the force. In general, elastomers deform when force is applied, but then return to their original shape when the force is removed. The elasticity exhibited by elastomeric materials may be characterized by a Young's modulus. Elastomeric materials having a Young's modulus of between about 1 Pa-1 TPa, more preferably between about 10 Pa-100 GPa, more preferably between about 20 Pa-1 GPa, more preferably between about 50 Pa-10 MPa, and more preferably between about 100 Pa-1 MPa are useful in accordance with the present invention, although elastomeric materials having a Young's modulus outside of these ranges could also be utilized depending upon the needs of a particular application.

The systems of the present invention may be fabricated from a wide variety of elastomers. In an exemplary aspect, the elastomeric layers may preferably be fabricated from silicone rubber. However, other suitable elastomers may also be used.

In an exemplary aspect of the present invention, the present systems are fabricated from an elastomeric polymer such as GE RTV 615 (formulation), a vinyl-silane crosslinked (type) silicone elastomer (family). However, the present systems are not limited to this one formulation, type or even this family of polymer; rather, nearly any elastomeric polymer is suitable. An important requirement for the preferred method of fabrication of the present microvalves is the ability to bond multiple layers of elastomers together. In the case of multilayer soft lithography, layers of elastomer are cured separately and then bonded together. This scheme requires that cured layers possess sufficient reactivity to bond together. Either the layers may be of the same type, and are capable of bonding to themselves, or they may be of two different types, and are capable of bonding to each other. Other possibilities include the use an adhesive between layers and the use of thermoset elastomers.

Given the tremendous diversity of polymer chemistries, precursors, synthetic methods, reaction conditions, and potential additives, there are a huge number of possible elastomer systems that could be used to make monolithic elastomeric microvalves and pumps. Variations in the materials used will most likely be driven by the need for particular material properties, i.e. solvent resistance, stiffness, gas permeability, or temperature stability.

There are many, many types of elastomeric polymers. A brief description of the most common classes of elastomers is presented here, with the intent of showing that even with relatively “standard” polymers, many possibilities for bonding exist. Common elastomeric polymers include polyisoprene, polybutadiene, polychloroprene, polyisobutylene, poly(styrene-butadiene-styrene), the polyurethanes, and silicones.

Polyisoprene, polybutadiene, polychloroprene:

Polyisoprene, polybutadiene, and polychloroprene are all polymerized from diene monomers, and therefore have one double bond per monomer when polymerized. This double bond allows the polymers to be converted to elastomers by vulcanization (essentially, sulfur is used to form crosslinks between the double bonds by heating). This would easily allow homogeneous multilayer soft lithography by incomplete vulcanization of the layers to be bonded; photoresist encapsulation would be possible by a similar mechanism.

Polyisobutylene:

Pure Polyisobutylene has no double bonds, but is crosslinked to use as an elastomer by including a small amount (˜1%) of isoprene in the polymerization. The isoprene monomers give pendant double bonds on the Polyisobutylene backbone, which may then be vulcanized as above.

Poly(styrene-butadiene-styrene):

Poly(styrene-butadiene-styrene) is produced by living anionic polymerization (that is, there is no natural chain-terminating step in the reaction), so “live” polymer ends can exist in the cured polymer. This makes it a natural candidate for the present photoresist encapsulation system (where there will be plenty of unreacted monomer in the liquid layer poured on top of the cured layer). Incomplete curing would allow homogeneous multilayer soft lithography (A to A bonding). The chemistry also facilitates making one layer with extra butadiene (“A”) and coupling agent and the other layer (“B”) with a butadiene deficit (for heterogeneous multilayer soft lithography). SBS is a “thermoset elastomer”, meaning that above a certain temperature it melts and becomes plastic (as opposed to elastic); reducing the temperature yields the elastomer again. Thus, layers can be bonded together by heating.

Polyurethanes:

Polyurethanes are produced from di-isocyanates (A-A) and di-alcohols or di-amines (B-B); since there are a large variety of di-isocyanates and di-alcohols/amines, the number of different types of polyurethanes is huge. The A vs. B nature of the polymers, however, would make them useful for heterogeneous multilayer soft lithography just as RTV 615 is: by using excess A-A in one layer and excess B-B in the other layer.

Silicones:

Silicone polymers probably have the greatest structural variety, and almost certainly have the greatest number of commercially available formulations. The vinyl-to-(Si—H) crosslinking of RTV 615 (which allows both heterogeneous multilayer soft lithography and photoresist encapsulation) has already been discussed, but this is only one of several crosslinking methods used in silicone polymer chemistry.

5. Operation of Device

FIGS. 7B and 7H together show the closing of a first flow channel by pressurizing a second flow channel, with FIG. 7B (a front sectional view cutting through flow channel 32 in corresponding FIG. 7A), showing an open first flow channel 30; with FIG. 7H showing first flow channel 30 closed by pressurization of the second flow channel 32.

Referring to FIG. 7B, first flow channel 30 and second flow channel 32 are shown. Membrane 25 separates the flow channels, forming the top of first flow channel 30 and the bottom of second flow channel 32. As can be seen, flow channel 30 is “open”.

As can be seen in FIG. 7H, pressurization of flow channel 32 (either by gas or liquid introduced therein) causes membrane 25 to deflect downward, thereby pinching off flow F passing through flow channel 30. Accordingly, by varying the pressure in channel 32, a linearly actuable valving system is provided such that flow channel 30 can be opened or closed by moving membrane 25 as desired. (For illustration purposes only, channel 30 in FIG. 7G is shown in a “mostly closed” position, rather than a “fully closed” position).

Since such valves are actuated by moving the roof of the channels themselves (i.e.: moving membrane 25) valves and pumps produced by this technique have a truly zero dead volume, and switching valves made by this technique have a dead volume approximately equal to the active volume of the valve, for example about 100×100×10 μm=100 pL. Such dead volumes and areas consumed by the moving membrane are approximately two orders of magnitude smaller than known conventional microvalves. Smaller and larger valves and switching valves are contemplated in the present invention, and a non-exclusive list of ranges of dead volume includes 1 aL to 1 uL, 100 aL to 100 nL, 1 fL to 10 nL, 100 fL to 1 nL, and 1 pL to 100 pL.

The extremely small volumes capable of being delivered by pumps and valves in accordance with the present invention represent a substantial advantage. Specifically, the smallest known volumes of fluid capable of being manually metered is around 0.1 μl. The smallest known volumes capable of being metered by automated systems is about ten-times larger (1 μl). Utilizing pumps and valves in accordance with the present invention, volumes of liquid of 10 nl or smaller can routinely be metered and dispensed. The accurate metering of extremely small volumes of fluid enabled by the present invention would be extremely valuable in a large number of biological applications, including diagnostic tests and assays.

Equation 1 represents a highly simplified mathematical model of deflection of a rectangular, linear, elastic, isotropic plate of uniform thickness by an applied pressure: w=(BPb ⁴)/(Eh ³), where:  (1)

-   -   w=deflection of plate;     -   B=shape coefficient (dependent upon length vs. width and support         of edges of plate);     -   P=applied pressure;     -   b=plate width     -   E=Young's modulus; and     -   h=plate thickness.

Thus even in this extremely simplified expression, deflection of an elastomeric membrane in response to a pressure will be a function of: the length, width, and thickness of the membrane, the flexibility of the membrane (Young's modulus), and the applied actuation force. Because each of these parameters will vary widely depending upon the actual dimensions and physical composition of a particular elastomeric device in accordance with the present invention, a wide range of membrane thicknesses and elasticity's, channel widths, and actuation forces are contemplated by the present invention.

It should be understood that the formula just presented is only an approximation, since in general the membrane does not have uniform thickness, the membrane thickness is not necessarily small compared to the length and width, and the deflection is not necessarily small compared to length, width, or thickness of the membrane. Nevertheless, the equation serves as a useful guide for adjusting variable parameters to achieve a desired response of deflection versus applied force.

FIGS. 8A and 8B illustrate valve opening vs. applied pressure for a 100 μm wide first flow channel 30 and a 50 μm wide second flow channel 32. The membrane of this device was formed by a layer of General Electric Silicones RTV 615 having a thickness of approximately 30 μm and a Young's modulus of approximately 750 kPa. FIGS. 21A and 21B show the extent of opening of the valve to be substantially linear over most of the range of applied pressures.

Air pressure was applied to actuate the membrane of the device through a 10 cm long piece of plastic tubing having an outer diameter of 0.025″ connected to a 25 mm piece of stainless steel hypodermic tubing with an outer diameter of 0.025″ and an inner diameter of 0.013″. This tubing was placed into contact with the control channel by insertion into the elastomeric block in a direction normal to the control channel. Air pressure was applied to the hypodermic tubing from an external LHDA miniature solenoid valve manufactured by Lee Co.

While control of the flow of material through the device has so far been described utilizing applied gas pressure, other fluids could be used.

For example, air is compressible, and thus experiences some finite delay between the time of application of pressure by the external solenoid valve and the time that this pressure is experienced by the membrane. In an alternative embodiment of the present invention, pressure could be applied from an external source to a noncompressible fluid such as water or hydraulic oils, resulting in a near-instantaneous transfer of applied pressure to the membrane. However, if the displaced volume of the valve is large or the control channel is narrow, higher viscosity of a control fluid may contribute to delay in actuation. The optimal medium for transferring pressure will therefore depend upon the particular application and device configuration, and both gaseous and liquid media are contemplated by the invention.

While external applied pressure as described above has been applied by a pump/tank system through a pressure regulator and external miniature valve, other methods of applying external pressure are also contemplated in the present invention, including gas tanks, compressors, piston systems, and columns of liquid. Also contemplated is the use of naturally occurring pressure sources such as may be found inside living organisms, such as blood pressure, gastric pressure, the pressure present in the cerebro-spinal fluid, pressure present in the intra-ocular space, and the pressure exerted by muscles during normal flexure. Other methods of regulating external pressure are also contemplated, such as miniature valves, pumps, macroscopic peristaltic pumps, pinch valves, and other types of fluid regulating equipment such as is known in the art.

As can be seen, the response of valves in accordance with embodiments of the present invention have been experimentally shown to be almost perfectly linear over a large portion of its range of travel, with minimal hysteresis. Accordingly, the present valves are ideally suited for microfluidic metering and fluid control. The linearity of the valve response demonstrates that the individual valves are well modeled as Hooke's Law springs. Furthermore, high pressures in the flow channel (i.e.: back pressure) can be countered simply by increasing the actuation pressure. Experimentally, the present inventors have achieved valve closure at back pressures of 70 kPa, but higher pressures are also contemplated. The following is a nonexclusive list of pressure ranges encompassed by the present invention: 10 Pa-25 MPa; 100 Pa-10 Mpa, 1 kPa-1 MPa, 1 kPa-300 kPa, 5 kPa-200 kPa, and 15 kPa-100 kPa.

While valves and pumps do not require linear actuation to open and close, linear response does allow valves to more easily be used as metering devices. In one embodiment of the invention, the opening of the valve is used to control flow rate by being partially actuated to a known degree of closure. Linear valve actuation makes it easier to determine the amount of actuation force required to close the valve to a desired degree of closure. Another benefit of linear actuation is that the force required for valve actuation may be easily determined from the pressure in the flow channel. If actuation is linear, increased pressure in the flow channel may be countered by adding the same pressure (force per unit area) to the actuated portion of the valve.

Linearity of a valve depends on the structure, composition, and method of actuation of the valve structure. Furthermore, whether linearity is a desirable characteristic in a valve depends on the application. Therefore, both linearly and non-linearly actuable valves are contemplated in the present invention, and the pressure ranges over which a valve is linearly actuable will vary with the specific embodiment.

FIG. 9 illustrates time response (i.e.: closure of valve as a function of time in response to a change in applied pressure) of a 100 μm×100 μm×10 μm RTV microvalve with 10-cm-long air tubing connected from the chip to a pneumatic valve as described above.

Two periods of digital control signal, actual air pressure at the end of the tubing and valve opening are shown in FIG. 9. The pressure applied on the control line is 100 kPa, which is substantially higher than the ˜40 kPa required to close the valve. Thus, when closing, the valve is pushed closed with a pressure 60 kPa greater than required. When opening, however, the valve is driven back to its rest position only by its own spring force (≦40 kPa). Thus, τ_(close) is expected to be smaller than τ_(open). There is also a lag between the control signal and control pressure response, due to the limitations of the miniature valve used to control the pressure. Calling such lags t and the 1/e time constants τ, the values are: t_(open)=3.63 ms, τ_(open)=1.88 ms, t_(close)=2.15 ms, τ_(dose)=0.51 ms. If 3τ each are allowed for opening and closing, the valve runs comfortably at 75 Hz when filled with aqueous solution.

If one used another actuation method which did not suffer from opening and closing lag, this valve would run at ˜375 Hz. Note also that the spring constant can be adjusted by changing the membrane thickness; this allows optimization for either fast opening or fast closing. The spring constant could also be adjusted by changing the elasticity (Young's modulus) of the membrane, as is possible by introducing dopant into the membrane or by utilizing a different elastomeric material to serve as the membrane (described above in conjunction with FIGS. 7C-H).

When experimentally measuring the valve properties as illustrated in FIG. 9 the valve opening was measured by fluorescence. In these experiments, the flow channel was filled with a solution of fluorescein isothiocyanate (FITC) in buffer (pH 8) and the fluorescence of a square area occupying the center ˜⅓rd of the channel is monitored on an epi-fluorescence microscope with a photomultiplier tube with a 10 kHz bandwidth. The pressure was monitored with a Wheatstone-bridge pressure sensor (SenSym SCC15GD2) pressurized simultaneously with the control line through nearly identical pneumatic connections.

6. Flow Channel Cross Sections

The flow channels of the present invention may optionally be designed with different cross sectional sizes and shapes, offering different advantages, depending upon their desired application. For example, the cross sectional shape of the lower flow channel may have a curved upper surface, either along its entire length or in the region disposed under an upper cross channel). Such a curved upper surface facilitates valve sealing, as follows.

Referring to FIG. 10, a cross sectional view (similar to that of FIG. 7B) through flow channels 30 and 32 is shown. As can be seen, flow channel 30 is rectangular in cross sectional shape. In an alternate preferred aspect of the invention, as shown in FIG. 10, the cross-section of a flow channel 30 instead has an upper curved surface.

Referring first to FIG. 10, when flow channel 32 is pressurized, the membrane portion 25 of elastomeric block 24 separating flow channels 30 and 32 will move downwardly to the successive positions shown by the dotted lines 25A, 25B, 25C, 25D, and 25E. As can be seen, incomplete sealing may possibly result at the edges of flow channel 30 adjacent planar substrate 14.

In the alternate preferred embodiment of FIG. 11, flow channel 30 a has a curved upper wall 25A. When flow channel 32 is pressurized, membrane portion 25 will move downwardly to the successive positions shown by dotted lines 25A2, 25A3, 25A4 and 25A5, with edge portions of the membrane moving first into the flow channel, followed by top membrane portions. An advantage of having such a curved upper surface at membrane 25A is that a more complete seal will be provided when flow channel 32 is pressurized. Specifically, the upper wall of the flow channel 30 will provide a continuous contacting edge against planar substrate 14, thereby avoiding the “island” of contact seen between wall 25 and the bottom of flow channel 30 in FIG. 10.

Another advantage of having a curved upper flow channel surface at membrane 25A is that the membrane can more readily conform to the shape and volume of the flow channel in response to actuation. Specifically, where a rectangular flow channel is employed, the entire perimeter (2× flow channel height, plus the flow channel width) must be forced into the flow channel. However where an arched flow channel is used, a smaller perimeter of material (only the semi-circular arched portion) must be forced into the channel. In this manner, the membrane requires less change in perimeter for actuation and is therefore more responsive to an applied actuation force to block the flow channel.

In an alternate aspect, (not illustrated), the bottom of flow channel 30 is rounded such that its curved surface mates with the curved upper wall 25A as seen in FIG. 20 described above.

In summary, the actual conformational change experienced by the membrane upon actuation will depend upon the configuration of the particular elastomeric structure. Specifically, the conformational change will depend upon the length, width, and thickness profile of the membrane, its attachment to the remainder of the structure, and the height, width, and shape of the flow and control channels and the material properties of the elastomer used. The conformational change may also depend upon the method of actuation, as actuation of the membrane in response to an applied pressure will vary somewhat from actuation in response to a magnetic or electrostatic force.

Moreover, the desired conformational change in the membrane will also vary depending upon the particular application for the elastomeric structure. In the simplest embodiments described above, the valve may either be open or closed, with metering to control the degree of closure of the valve. In other embodiments however, it may be desirable to alter the shape of the membrane and/or the flow channel in order to achieve more complex flow regulation. For instance, the flow channel could be provided with raised protrusions beneath the membrane portion, such that upon actuation the membrane shuts off only a percentage of the flow through the flow channel, with the percentage of flow blocked insensitive to the applied actuation force.

Many membrane thickness profiles and flow channel cross-sections are contemplated by the present invention, including rectangular, trapezoidal, circular, ellipsoidal, parabolic, hyperbolic, and polygonal, as well as sections of the above shapes. More complex cross-sectional shapes, such as the embodiment with protrusions discussed immediately above or an embodiment having concavities in the flow channel, are also contemplated by the present invention.

In addition, while the invention is described primarily above in conjunction with an embodiment wherein the walls and ceiling of the flow channel are formed from elastomer, and the floor of the channel is formed from an underlying substrate, the present invention is not limited to this particular orientation. Walls and floors of channels could also be formed in the underlying substrate, with only the ceiling of the flow channel constructed from elastomer. This elastomer flow channel ceiling would project downward into the channel in response to an applied actuation force, thereby controlling the flow of material through the flow channel. In general, monolithic elastomer structures as described elsewhere in the instant application are preferred for microfluidic applications. However, it may be useful to employ channels formed in the substrate where such an arrangement provides advantages. For instance, a substrate including optical waveguides could be constructed so that the optical waveguides direct light specifically to the side of a microfluidic channel.

7. Alternate Valve Actuation Techniques

In addition to pressure based actuation systems described above, optional electrostatic and magnetic actuation systems are also contemplated, as follows.

Electrostatic actuation can be accomplished by forming oppositely charged electrodes (which will tend to attract one another when a voltage differential is applied to them) directly into the monolithic elastomeric structure. For example, referring to FIG. 7B, an optional first electrode 70 (shown in phantom) can be positioned on (or in) membrane 25 and an optional second electrode 72 (also shown in phantom) can be positioned on (or in) planar substrate 14. When electrodes 70 and 72 are charged with opposite polarities, an attractive force between the two electrodes will cause membrane 25 to deflect downwardly, thereby closing the “valve” (i.e.: closing flow channel 30).

For the membrane electrode to be sufficiently conductive to support electrostatic actuation, but not so mechanically stiff so as to impede the valve's motion, a sufficiently flexible electrode must be provided in or over membrane 25. Such an electrode may be provided by a thin metallization layer, doping the polymer with conductive material, or making the surface layer out of a conductive material.

In an exemplary aspect, the electrode present at the deflecting membrane can be provided by a thin metallization layer which can be provided, for example, by sputtering a thin layer of metal such as 20 nm of gold. In addition to the formation of a metallized membrane by sputtering, other metallization approaches such as chemical epitaxy, evaporation, electroplating, and electroless plating are also available. Physical transfer of a metal layer to the surface of the elastomer is also available, for example by evaporating a metal onto a flat substrate to which it adheres poorly, and then placing the elastomer onto the metal and peeling the metal off of the substrate.

A conductive electrode 70 may also be formed by depositing carbon black (i.e. Cabot Vulcan XC72R) on the elastomer surface, either by wiping on the dry powder or by exposing the elastomer to a suspension of carbon black in a solvent which causes swelling of the elastomer, (such as a chlorinated solvent in the case of PDMS). Alternatively, the electrode 70 may be formed by constructing the entire layer 20 out of elastomer doped with conductive material (i.e. carbon black or finely divided metal particles). Yet further alternatively, the electrode may be formed by electrostatic deposition, or by a chemical reaction that produces carbon. In experiments conducted by the present inventors, conductivity was shown to increase with carbon black concentration from 5.6×10⁻¹⁶ to about 5×10⁻³ (Ω-cm)⁻¹. The lower electrode 72, which is not required to move, may be either a compliant electrode as described above, or a conventional electrode such as evaporated gold, a metal plate, or a doped semiconductor electrode.

Magnetic actuation of the flow channels can be achieved by fabricating the membrane separating the flow channels with a magnetically polarizable material such as iron, or a permanently magnetized material such as polarized NdFeB. In experiments conducted by the present inventors, magnetic silicone was created by the addition of iron powder (about 1 um particle size), up to 20% iron by weight.

Where the membrane is fabricated with a magnetically polarizable material, the membrane can be actuated by attraction in response to an applied magnetic field Where the membrane is fabricated with a material capable of maintaining permanent magnetization, the material can first be magnetized by exposure to a sufficiently high magnetic field, and then actuated either by attraction or repulsion in response to the polarity of an applied inhomogeneous magnetic field.

The magnetic field causing actuation of the membrane can be generated in a variety of ways. In one embodiment, the magnetic field is generated by an extremely small inductive coil formed in or proximate to the elastomer membrane. The actuation effect of such a magnetic coil would be localized, allowing actuation of individual pump and/or valve structures. Alternatively, the magnetic field could be generated by a larger, more powerful source, in which case actuation would be global and would actuate multiple pump and/or valve structures at one time.

It is also possible to actuate the device by causing a fluid flow in the control channel based upon the application of thermal energy, either by thermal expansion or by production of gas from liquid. For example, in one alternative embodiment in accordance with the present invention, a pocket of fluid (e.g. in a fluid-filled control channel) is positioned over the flow channel. Fluid in the pocket can be in communication with a temperature variation system, for example a heater. Thermal expansion of the fluid, or conversion of material from the liquid to the gas phase, could result in an increase in pressure, closing the adjacent flow channel. Subsequent cooling of the fluid would relieve pressure and permit the flow channel to open.

8. Networked Systems

FIGS. 12A and 12B show a views of a single on/off valve, identical to the systems set forth above, (for example in FIG. 7A). FIGS. 13A and 13B shows a peristaltic pumping system comprised of a plurality of the single addressable on/off valves as seen in FIG. 12, but networked together. FIG. 14 is a graph showing experimentally achieved pumping rates vs. frequency for the peristaltic pumping system of FIG. 13. FIGS. 15A and 15B show a schematic view of a plurality of flow channels which are controllable by a single control line. This system is also comprised of a plurality of the single addressable on/off valves of FIG. 12, multiplexed together, but in a different arrangement than that of FIG. 12. FIG. 16 is a schematic illustration of a multiplexing system adapted to permit fluid flow through selected channels, comprised of a plurality of the single on/off valves of FIG. 12, joined or networked together.

Referring first to FIGS. 12A and 12B, a schematic of flow channels 30 and 32 is shown. Flow channel 30 preferably has a fluid (or gas) flow F passing therethrough. Flow channel 32, (which crosses over flow channel 30, as was already explained herein), is pressurized such that membrane 25 separating the flow channels may be depressed into the path of flow channel 30, shutting off the passage of flow F therethrough, as has been explained. As such, “flow channel” 32 can also be referred to as a “control line” which actuates a single valve in flow channel 30. In FIGS. 12 to 15, a plurality of such addressable valves are joined or networked together in various arrangements to produce pumps, capable of peristaltic pumping, and other fluidic logic applications.

Referring to FIGS. 13A and 13B, a system for peristaltic pumping is provided, as follows. A flow channel 30 has a plurality of generally parallel flow channels (i.e.: control lines) 32A, 32B and 32C passing thereover. By pressurizing control line 32A, flow F through flow channel 30 is shut off under membrane 25A at the intersection of control line 32A and flow channel 30. Similarly, (but not shown), by pressurizing control line 32B, flow F through flow channel 30 is shut off under membrane 25B at the intersection of control line 32B and flow channel 30, etc.

Each of control lines 32A, 32B, and 32C is separately addressable. Therefore, peristalsis may be actuated by the pattern of actuating 32A and 32C together, followed by 32A, followed by 32A and 32B together, followed by 32B, followed by 32B and C together, etc. This corresponds to a successive “101, 100, 110, 010, 011, 001” pattern, where “0” indicates “valve open” and “1” indicates “valve closed.” This peristaltic pattern is also known as a 120° pattern (referring to the phase angle of actuation between three valves). Other peristaltic patterns are equally possible, including 60° and 90° patterns.

In experiments performed by the inventors, a pumping rate of 2.35 nL/s was measured by measuring the distance traveled by a column of water in thin (0.5 mm i.d.) tubing; with 100×100×10 μm valves under an actuation pressure of 40 kPa. The pumping rate increased with actuation frequency until approximately 75 Hz, and then was nearly constant until above 200 Hz. The valves and pumps are also quite durable and the elastomer membrane, control channels, or bond have never been observed to fail. In experiments performed by the inventors, none of the valves in the peristaltic pump described herein show any sign of wear or fatigue after more than 4 million actuations. In addition to their durability, they are also gentle. A solution of E. Coli pumped through a channel and tested for viability showed a 94% survival rate.

FIG. 14 is a graph showing experimentally achieved pumping rates vs. frequency for the peristaltic pumping system of FIG. 13.

FIGS. 15A and 15B illustrates another way of assembling a plurality of the addressable valves of FIG. 12. Specifically, a plurality of parallel flow channels 30A, 30B, and 30C are provided. Flow channel (i.e.: control line) 32 passes thereover across flow channels 30A, 30B, and 30C. Pressurization of control line 32 simultaneously shuts off flows F1, F2 and F3 by depressing membranes 25A, 25B, and 25C located at the intersections of control line 32 and flow channels 30A, 30B, and 30C.

FIG. 16 is a schematic illustration of a multiplexing system adapted to selectively permit fluid to flow through selected channels, as follows. The downward deflection of membranes separating the respective flow channels from a control line passing there above (for example, membranes 25A, 25B, and 25C in FIGS. 15A and 15B) depends strongly upon the membrane dimensions. Accordingly, by varying the widths of flow channel control line 32 in FIGS. 15A and 15B, it is possible to have a control line pass over multiple flow channels, yet only actuate (i.e.: seal) desired flow channels. FIG. 16 illustrates a schematic of such a system, as follows.

A plurality of parallel flow channels 30A, 30B, 30C, 30D, 30E and 30F are positioned under a plurality of parallel control lines 32A, 32B, 32C, 32D, 32E and 32F. Control channels 32A, 32B, 32C, 32D, 32E and 32F are adapted to shut off fluid flows F1, F2, F3, F4, F5 and F6 passing through parallel flow channels 30A, 30B, 30C, 30D, 30E and 30F using any of the valving systems described above, with the following modification.

Each of control lines 32A, 32B, 32C, 32D, 32E and 32F have both wide and narrow portions. For example, control line 32A is wide in locations disposed over flow channels 30A, 30C and 30E. Similarly, control line 32B is wide in locations disposed over flow channels 30B, 30D and 30F, and control line 32C is wide in locations disposed over flow channels 30A, 30B, 30E and 30F.

At the locations where the respective control line is wide, its pressurization will cause the membrane (25) separating the flow channel and the control line to depress significantly into the flow channel, thereby blocking the flow passage therethrough. Conversely, in the locations where the respective control line is narrow, membrane (25) will also be narrow. Accordingly, the same degree of pressurization will not result in membrane (25) becoming depressed into the flow channel (30). Therefore, fluid passage there under will not be blocked.

For example, when control line 32A is pressurized, it will block flows F1, F3 and F5 in flow channels 30A, 30C and 30E. Similarly, when control line 32C is pressurized, it will block flows F1, F2, F5 and F6 in flow channels 30A, 30B, 30E and 30F. As can be appreciated, more than one control line can be actuated at the same time. For example, control lines 32A and 32C can be pressurized simultaneously to block all fluid flow except F4 (with 32A blocking F1, F3 and F5; and 32C blocking F1, F2, F5 and F6).

By selectively pressurizing different control lines (32) both together and in various sequences, a great degree of fluid flow control can be achieved. Moreover, by extending the present system to more than six parallel flow channels (30) and more than four parallel control lines (32), and by varying the positioning of the wide and narrow regions of the control lines, very complex fluid flow control systems may be fabricated. A property of such systems is that it is possible to turn on any one flow channel out of n flow channels with only 2(log₂n) control lines.

9. Selectively Addressable Reaction Chambers Along Flow Lines

In a further embodiment of the invention, illustrated in FIGS. 17A-D, a system for selectively directing fluid flow into one more of a plurality of reaction chambers disposed along a flow line is provided.

FIG. 17A shows a top view of a flow channel 30 having a plurality of reaction chambers 80A and 80B disposed there along. Preferably flow channel 30 and reaction chambers 80A and 80B are formed together as recesses into the bottom surface of a first layer 100 of elastomer.

FIG. 17B shows a bottom plan view of another elastomeric layer 110 with two control lines 32A and 32B each being generally narrow, but having wide extending portions 33A and 33B formed as recesses therein.

As seen in the exploded view of FIG. 17C, and assembled view of FIG. 17D, elastomeric layer 110 is placed over elastomeric layer 100. Layers 100 and 110 are then bonded together, and the integrated system operates to selectively direct fluid flow F (through flow channel 30) into either or both of reaction chambers 80A and 80B, as follows. Pressurization of control line 32A will cause the membrane 25 (i.e.: the thin portion of elastomer layer 100 located below extending portion 33A and over regions 82A of reaction chamber 80A) to become depressed, thereby shutting off fluid flow passage in regions 82A, effectively sealing reaction chamber 80 from flow channel 30. As can also be seen, extending portion 33A is wider than the remainder of control line 32A. As such, pressurization of control line 32A will not result in control line 32A sealing flow channel 30.

As can be appreciated, either or both of control lines 32A and 32B can be actuated at once. When both control lines 32A and 32B are pressurized together, sample flow in flow channel 30 will enter neither of reaction chambers 80A or 80B.

The concept of selectably controlling fluid introduction into various addressable reaction chambers disposed along a flow line (FIGS. 17A-D) can be combined with concept of selectably controlling fluid flow through one or more of a plurality of parallel flow lines (FIG. 16) to yield a system in which a fluid sample or samples can be can be sent to any particular reaction chamber in an array of reaction chambers. An example of such a system is provided in FIG. 18, in which parallel control channels 32A, 32B and 32C with extending portions 34 (all shown in phantom) selectively direct fluid flows F1 and F2 into any of the array of reaction wells 80A, 80B, 80C or 80D as explained above; while pressurization of control lines 32C and 32D selectively shuts off flows F2 and F1, respectively.

In yet another novel embodiment, fluid passage between parallel flow channels is possible. Referring to FIG. 19, either or both of control lines 32A or 32D can be depressurized such that fluid flow through lateral passageways 35 (between parallel flow channels 30A and 30B) is permitted. In this aspect of the invention, pressurization of control lines 32C and 32D would shut flow channel 30A between 35A and 35B, and would also shut lateral passageways 35B. As such, flow entering as flow F1 would sequentially travel through 30A, 35A and leave 30B as flow F4.

10. Switchable Flow Arrays

In yet another novel embodiment, fluid passage can be selectively directed to flow in either of two perpendicular directions. An example of such a “switchable flow array” system is provided in FIGS. 20A-D. FIG. 20A shows a bottom view of a first layer of elastomer 90, (or any other suitable substrate), having a bottom surface with a pattern of recesses forming a flow channel grid defined by an array of solid posts 92, each having flow channels passing there around.

In preferred aspects, an additional layer of elastomer is bound to the top surface of layer 90 such that fluid flow can be selectively directed to move either in direction F1, or perpendicular direction F2. FIG. 20 is a bottom view of the bottom surface of the second layer of elastomer 95 showing recesses formed in the shape of alternating “vertical” control lines 96 and “horizontal” control lines 94. “Vertical” control lines 96 have the same width there along, whereas “horizontal” control lines 94 have alternating wide and narrow portions, as shown.

Elastomeric layer 95 is positioned over top of elastomeric layer 90 such that “vertical” control lines 96 are positioned over posts 92 as shown in FIG. 20C and “horizontal” control lines 94 are positioned with their wide portions between posts 92, as shown in FIG. 20D.

As can be seen in FIG. 20C, when “vertical” control lines 96 are pressurized, the membrane of the integrated structure formed by the elastomeric layer initially positioned between layers 90 and 95 in regions 98 will be deflected downwardly over the array of flow channels such that flow in only able to pass in flow direction F2 (i.e.: vertically), as shown.

As can be seen in FIG. 20D, when “horizontal” control lines 94 are pressurized, the membrane of the integrated structure formed by the elastomeric layer initially positioned between layers 90 and 95 in regions 99 will be deflected downwardly over the array of flow channels, (but only in the regions where they are widest), such that flow in only able to pass in flow direction F1 (i.e.: horizontally), as shown.

The design illustrated in FIG. 20 allows a switchable flow array to be constructed from only two elastomeric layers, with no vertical vias passing between control lines in different elastomeric layers required. If all vertical flow control lines 94 are connected, they may be pressurized from one input. The same is true for all horizontal flow control lines 96.

11. Normally-Closed Valve Structure

FIGS. 7B and 7H above depict a valve structure in which the elastomeric membrane is moveable from a first relaxed position to a second actuated position in which the flow channel is blocked. However, the present invention is not limited to this particular valve configuration.

FIGS. 21A-J show a variety of views of a normally-closed valve structure in which the elastomeric membrane is moveable from a first relaxed position blocking a flow channel, to a second actuated position in which the flow channel is open, utilizing a negative control pressure.

FIG. 21A shows a plan view, and FIG. 21B shows a cross sectional view along line 42B-42B′, of normally-closed valve 4200 in an unactuated state. Flow channel 4202 and control channel 4204 are formed in elastomeric block 4206 overlying substrate 4205. Flow channel 4202 includes a first portion 4202 a and a second portion 4202 b separated by separating portion 4208. Control channel 4204 overlies separating portion 4208. As shown in FIG. 21B, in its relaxed, unactuated position, separating portion 4208 remains positioned between flow channel portions 4202 a and 4202 b, interrupting flow channel 4202.

FIG. 21C shows a cross-sectional view of valve 4200 wherein separating portion 4208 is in an actuated position. When the pressure within control channel 4204 is reduced to below the pressure in the flow channel (for example by vacuum pump), separating portion 4208 experiences an actuating force drawing it into control channel 4204. As a result of this actuation force membrane 4208 projects into control channel 4204, thereby removing the obstacle to a flow of material through flow channel 4202 and creating a passageway 4203. Upon elevation of pressure within control channel 4204, separating portion 4208 will assume its natural position, relaxing back into and obstructing flow channel 4202.

The behavior of the membrane in response to an actuation force may be changed by varying the width of the overlying control channel. Accordingly, FIGS. 21D-42H show plan and cross-sectional views of an alternative embodiment of a normally-closed valve 4201 in which control channel 4207 is substantially wider than separating portion 4208. As shown in cross-sectional views FIGS. 21E-F along line 42E-42E′ of FIG. 21D, because a larger area of elastomeric material is required to be moved during actuation, the actuation force necessary to be applied is reduced.

FIGS. 21G and 21H show a cross-sectional views along line 40G-40G′ of FIG. 21D. In comparison with the unactuated valve configuration shown in FIG. 21G, FIG. 21H shows that reduced pressure within wider control channel 4207 may under certain circumstances have the unwanted effect of pulling underlying elastomer 4206 away from substrate 4205, thereby creating undesirable void 4212.

Accordingly, FIG. 21I shows a plan view, and FIG. 21J shows a cross-sectional view along line 21J-21F of FIG. 21I, of valve structure 4220 which avoids this problem by featuring control line 4204 with a minimum width except in segment 4204 a overlapping separating portion 4208. As shown in FIG. 21J, even under actuated conditions the narrower cross-section of control channel 4204 reduces the attractive force on the underlying elastomer material 4206, thereby preventing this elastomer material from being drawn away from substrate 4205 and creating an undesirable void.

While a normally-closed valve structure actuated in response to pressure is shown in FIGS. 21A-21J, a normally-closed valve in accordance with the present invention is not limited to this configuration. For example, the separating portion obstructing the flow channel could alternatively be manipulated by electric or magnetic fields, as described extensively above.

12. Side-Actuated Valve

While the above description has focused upon microfabricated elastomeric valve structures in which a control channel is positioned above and separated by an intervening elastomeric membrane from an underlying flow channel, the present invention is not limited to this configuration. FIGS. 22A and 22B show plan views of one embodiment of a side-actuated valve structure in accordance with one embodiment of the present invention.

FIG. 22A shows side-actuated valve structure 4800 in an unactuated position. Flow channel 4802 is formed in elastomeric layer 4804. Control channel 4806 abutting flow channel 4802 is also formed in elastomeric layer 4804. Control channel 4806 is separated from flow channel 4802 by elastomeric membrane portion 4808. A second elastomeric layer (not shown) is bonded over bottom elastomeric layer 4804 to enclose flow channel 4802 and control channel 4806.

FIG. 22B shows side-actuated valve structure 4800 in an actuated position. In response to a build up of pressure within control channel 4806, membrane 4808 deforms into flow channel 4802, blocking flow channel 4802. Upon release of pressure within control channel 4806, membrane 4808 would relax back into control channel 4806 and open flow channel 4802.

While a side-actuated valve structure actuated in response to pressure is shown in FIGS. 22A and 22B, a side-actuated valve in accordance with the present invention is not limited to this configuration. For example, the elastomeric membrane portion located between the abutting flow and control channels could alternatively be manipulated by electric or magnetic fields, as described extensively above.

13. Composite Structures

Microfabricated elastomeric structures of the present invention may be combined with non-elastomeric materials to create composite structures. FIG. 23 shows a cross-sectional view of one embodiment of a composite structure in accordance with the present invention. FIG. 23 shows composite valve structure 5700 including first, thin elastomer layer 5702 overlying semiconductor-type substrate 5704 having channel 5706 formed therein. Second, thicker elastomer layer 5708 overlies first elastomer layer 5702. Actuation of first elastomer layer 5702 to drive it into channel 5706, will cause composite structure 5700 to operate as a valve.

FIG. 24 shows a cross-sectional view of a variation on this theme, wherein thin elastomer layer 5802 is sandwiched between two hard, semiconductor substrates 5804 and 5806, with lower substrate 5804 featuring channel 5808. Again, actuation of thin elastomer layer 5802 to drive it into channel 5808 will cause composite structure 5810 to operate as a valve.

The structures shown in FIG. 23 or 24 may be fabricated utilizing either the multilayer soft lithography or encapsulation techniques described above. In the multilayer soft lithography method, the elastomer layer(s) would be formed and then placed over the semiconductor substrate bearing the channel. In the encapsulation method, the channel would be first formed in the semiconductor substrate, and then the channel would be filled with a sacrificial material such as photoresist. The elastomer would then be formed in place over the substrate, with removal of the sacrificial material producing the channel overlaid by the elastomer membrane. As is discussed in detail below in connection with bonding of elastomer to other types of materials, the encapsulation approach may result in a stronger seal between the elastomer membrane component and the underlying nonelastomer substrate component.

As shown in FIGS. 23 and 24, a composite structure in accordance with embodiments of the present invention may include a hard substrate that bears a passive feature such as a channels. However, the present invention is not limited to this approach, and the underlying hard substrate may bear active features that interact with an elastomer component bearing a recess. This is shown in FIG. 25, wherein composite structure 5900 includes elastomer component 5902 containing recess 5904 having walls 5906 and ceiling 5908. Ceiling 5908 forms flexible membrane portion 5909. Elastomer component 5902 is sealed against substantially planar nonelastomeric component 5910 that includes active device 5912. Active device 5912 may interact with material present in recess 5904 and/or flexible membrane portion 5909.

Many Types of active structures may be present in the nonelastomer substrate. Active structures that could be present in an underlying hard substrate include, but are not limited to, resistors, capacitors, photodiodes, transistors, chemical field effect transistors (chem FET's), amperometric/coulometric electrochemical sensors, fiber optics, fiber optic interconnects, light emitting diodes, laser diodes, vertical cavity surface emitting lasers (VCSEL's), micromirrors, accelerometers, pressure sensors, flow sensors, CMOS imaging arrays, CCD cameras, electronic logic, microprocessors, thermistors, Peltier coolers, waveguides, resistive heaters, chemical sensors, strain gauges, inductors, actuators (including electrostatic, magnetic, electromagnetic, bimetallic, piezoelectric, shape-memory-alloy based, and others), coils, magnets, electromagnets, magnetic sensors (such as those used in hard drives, superconducting quantum interference devices (SQUIDS) and other types), radio frequency sources and receivers, microwave frequency sources and receivers, sources and receivers for other regions of the electromagnetic spectrum, radioactive particle counters, and electrometers.

As is well known in the art, a vast variety of technologies can be utilized to fabricate active features in semiconductor and other types of hard substrates, including but not limited printed circuit board (PCB) technology, CMOS, surface micromachining, bulk micromachining, printable polymer electronics, and TFT and other amorphous/polycrystalline techniques as are employed to fabricate laptop and flat screen displays.

A variety of approaches can be employed to seal the elastomeric structure against the nonelastomeric substrate, ranging from the creation of a Van der Waals bond between the elastomeric and nonelastomeric components, to creation of covalent or ionic bonds between the elastomeric and nonelastomeric components of the composite structure. Example approaches to sealing the components together are discussed below, approximately in order of increasing strength.

A first approach is to rely upon the simple hermetic seal resulting from Van der Waals bonds formed when a substantially planar elastomer layer is placed into contact with a substantially planar layer of a harder, non-elastomer material. In one embodiment, bonding of RTV elastomer to a glass substrate created a composite structure capable of withstanding up to about 3-4 psi of pressure. This may be sufficient for many potential applications.

A second approach is to utilize a liquid layer to assist in bonding. One example of this involves bonding elastomer to a hard glass substrate, wherein a weakly acidic solution (5 μl HCl in H₂O, pH 2) was applied to a glass substrate. The elastomer component was then placed into contact with the glass substrate, and the composite structure baked at 37° C. to remove the water. This resulted in a bond between elastomer and non-elastomer able to withstand a pressure of about 20 psi. In this case, the acid may neutralize silanol groups present on the glass surface, permitting the elastomer and non-elastomer to enter into good Van der Waals contact with each other.

Exposure to ethanol can also cause device components to adhere together. In one embodiment, an RTV elastomer material and a glass substrate were washed with ethanol and then dried under Nitrogen. The RTV elastomer was then placed into contact with the glass and the combination baked for 3 hours at 80° C. Optionally, the RTV may also be exposed to a vacuum to remove any air bubbles trapped between the slide and the RTV. The strength of the adhesion between elastomer and glass using this method has withstood pressures in excess of 35 psi. The adhesion created using this method is not permanent, and the elastomer may be peeled off of the glass, washed, and resealed against the glass. This ethanol washing approach can also be employed used to cause successive layers of elastomer to bond together with sufficient strength to resist a pressure of 30 psi. In alternative embodiments, chemicals such as other alcohols or diols could be used to promote adhesion between layers.

An embodiment of a method of promoting adhesion between layers of a microfabricated structure in accordance with the present invention comprises exposing a surface of a first component layer to a chemical, exposing a surface of a second component layer to the chemical, and placing the surface of the first component layer into contact with the surface of the second elastomer layer.

A third approach is to create a covalent chemical bond between the elastomer component and functional groups introduced onto the surface of a nonelastomer component. Examples of derivitization of a nonelastomer substrate surface to produce such functional groups include exposing a glass substrate to agents such as vinyl silane or aminopropyltriethoxy silane (APTES), which may be useful to allow bonding of the glass to silicone elastomer and polyurethane elastomer materials, respectively.

A fourth approach is to create a covalent chemical bond between the elastomer component and a functional group native to the surface of the nonelastomer component. For example, RTV elastomer can be created with an excess of vinyl groups on its surface. These vinyl groups can be caused to react with corresponding functional groups present on the exterior of a hard substrate material, for example the Si—H bonds prevalent on the surface of a single crystal silicon substrate after removal of native oxide by etching. In this example, the strength of the bond created between the elastomer component and the nonelastomer component has been observed to exceed the materials strength of the elastomer components.

14. Cell Pen

In yet a further application of the present invention, an elastomeric structure can be utilized to manipulate organisms or other biological material. FIGS. 26A-D show plan views of one embodiment of a cell pen structure in accordance with the present invention.

Cell pen array 4400 features an array of orthogonally-oriented flow channels 4402, with an enlarged “pen” structure 4404 at the intersection of alternating flow channels. Valve 4406 is positioned at the entrance and exit of each pen structure 4404. Peristaltic pump structures 4408 are positioned on each horizontal flow channel and on the vertical flow channels lacking a cell pen structure.

Cell pen array 4400 of FIG. 26A has been loaded with cells A-H that have been previously sorted. FIGS. 26B-C show the accessing and removal of individually stored cell C by 1) opening valves 4406 on either side of adjacent pens 4404 a and 4404 b, 2) pumping horizontal flow channel 4402 a to displace cells C and G, and then 3) pumping vertical flow channel 4402 b to remove cell C. FIG. 26D shows that second cell G is moved back into its prior position in cell pen array 4400 by reversing the direction of liquid flow through horizontal flow channel 4402 a.

The cross-flow channel architecture illustrated shown in FIGS. 26A-D can be used to perform functions other than the cell pen just described. For example, the cross-flow channel architecture can be utilized in mixing applications.

This is shown in FIGS. 28A-B, which illustrate a plan view of mixing steps performed by a microfabricated structure in accordance another embodiment of the present invention. Specifically, portion 7400 of a microfabricated mixing structure comprises first flow channel 7402 orthogonal to and intersecting with second flow channel 7404. Control channels 7406 overlie flow channels 7402 and 7404 and form valve pairs 7408 a-b and 7408 c-d that surround each intersection 7412.

As shown in FIG. 28A, valve pair 7408 a-b is initially opened while valve pair 7408 c-d is closed, and fluid sample 7410 is flowed to intersection 7412 through flow channel 7402. Valve pair 7408 c-d is then actuated, trapping fluid sample 7410 at intersection 7412.

Next, as shown in FIG. 28B, valve pairs 7408 a-b and 7408 c-d are opened, such that fluid sample 7410 is injected from intersection 7412 into flow channel 7404 bearing a cross-flow of fluid. The process shown in FIGS. 28A-B can be repeated to accurately dispense any number of fluid samples down cross-flow channel 7404. FIG. 54 plots Log(R/B) vs. number of slugs injected for one embodiment of a cross-flow injection system in accordance with the present invention. The reproducibility and relative independence of metering by cross-flow injection from process parameters such as flow resistance is further evidenced by FIG. 54, which plots injected volume versus number of injection cycles for cross-channel flow injection under a variety of flow conditions. FIG. 54 shows that volumes metered by cross-flow injection techniques increase on a linear basis over a succession of injection cycles. This linear relationship between volume and number of injection cycles is relatively independent of flow resistance parameters such as elevated fluid viscosity (imparted by adding 25% glycerol) and the length of the flow channel (1.0-2.5 cm).

While the embodiment shown and described above in connection with FIGS. 28A-B utilizes linked valve pairs on opposite sides of the flow channel intersections, this is not required by the present invention. Other configurations, including linking of adjacent valves of an intersection, or independent actuation of each valve surrounding an intersection, are possible to provide the desired flow characteristics. With the independent valve actuation approach however, it should be recognized that separate control structures would be utilized for each valve, complicating device layout.

15. Metering by Volume Exclusion

Many high throughput screening and diagnostic applications call for accurate combination and of different reagents in a reaction chamber. Given that it is frequently necessary to prime the channels of a microfluidic device in order to ensure fluid flow, it may be difficult to ensure mixed solutions do not become diluted or contaminated by the contents of the reaction chamber prior to sample introduction.

Volume exclusion is one technique enabling precise metering of the introduction of fluids into a reaction chamber. In this approach, a reaction chamber may be completely or partially emptied prior to sample injection. This method reduces contamination from residual contents of the chamber contents, and may be used to accurately meter the introduction of solutions in a reaction chamber.

Specifically, FIGS. 29A-D show cross-sectional views of a reaction chamber in which volume exclusion is employed to meter reactants. FIG. 29A shows a cross-sectional view of portion 6300 of a microfluidic device comprising first elastomer layer 6302 overlying second elastomer layer 6304. First elastomer layer 6302 includes control chamber 6306 in fluid communication with a control channel (not shown). Control chamber 6306 overlies and is separated from dead-end reaction chamber 6308 of second elastomer layer 6304 by membrane 6310. Second elastomer layer 6304 further comprises flow channel 6312 leading to dead-end reaction chamber 6308.

FIG. 29B shows the result of a pressure increase within control chamber 6306. Specifically, increased control chamber pressure causes membrane 6310 to flex downward into reaction chamber 6308, reducing by volume V the effective volume of reaction chamber 6308. This in turn excludes an equivalent volume V of reactant from reaction chamber 6308, such that volume V of first reactant X is output from flow channel 6312. The exact correlation between a pressure increase in control chamber 6306 and the volume of material output from flow channel 6312 can be precisely calibrated.

As shown in FIG. 29C, while elevated pressure is maintained within control chamber 6306, volume V′ of second reactant Y is placed into contact with flow channel 6312 and reaction chamber 6308.

In the next step shown in FIG. 29D, pressure within control chamber 6306 is reduced to original levels. As a result, membrane 6310 relaxes and the effective volume of reaction chamber 6308 increases. Volume V of second reactant Y is sucked into the device. By varying the relative size of the reaction and control chambers, it is possible to accurately mix solutions at a specified relative concentration. It is worth noting that the amount of the second reactant Y that is sucked into the device is solely dependent upon the excluded volume V, and is independent of volume V′ of Y made available at the opening of the flow channel.

While FIGS. 29A-D show a simple embodiment of the present invention involving a single reaction chamber, in more complex embodiments parallel structures of hundreds or thousands of reaction chambers could be actuated by a pressure increase in one control line.

Moreover, while the above description illustrates two reactants being combined at a relative concentration that fixed by the size of the control and reaction chambers, a volume exclusion technique could be employed to combine several reagents at variable concentrations in a single reaction chamber. One possible approach is to use several, separately addressable control chambers above each reaction chamber. An example of this architecture would be to have ten separate control lines instead of a single control chamber, allowing ten equivalent volumes to be pushed out or sucked in.

Another possible approach would utilize a single control chamber overlying the entire reaction chamber, with the effective volume of the reaction chamber modulated by varying the control chamber pressure. In this manner, analog control over the effective volume of the reaction chamber is possible. Analog volume control would in turn permit the combination of many solutions reactants at arbitrary relative concentrations.

An embodiment of a method of metering a volume of fluid in accordance with the present invention comprises providing a chamber having a volume in an elastomeric block separated from a control recess by an elastomeric membrane, and supplying a pressure to the control recess such that the membrane is deflected into the chamber and the volume is reduced by a calibrated amount, excluding from the chamber the calibrated volume of fluid.

II. Microfluidic Free Interface Diffusion Techniques

A microfluidic free interface (μFI) in accordance with embodiments of the present invention is a localized interface between at least one static fluid and another fluid wherein mixing between them is dominated by diffusion rather than by convective flow. For the purposes of this application, the term “fluid” refers to a material having a viscosity below a particular maximum. Examples of such maximum viscosities include but are not limited to 1000 CPoise, 900 CPoise, 800 CPoise, 700 CPoise, 600 CPoise, 500 CPoise, 400 CPoise, 300 CPoise, 250 CPoise, and 100 CPoise, and therefore exclude gels or polymers containing materials trapped therein.

In a microfluidic free interface in accordance with an embodiment of the present invention, at least one dimension of the interface is restricted in magnitude such that viscous forces dominate other forces. For example, in a microfluidic free interface in accordance with an embodiment of the present invention, the dominant forces acting upon the fluids are viscous rather than buoyant, and hence the microfluidic free interface may be characterized by an extremely low Grashof number (see discussion below). The microfluidic free interface may also be characterized by its localized nature relative to the total volumes of the fluids, such that the volumes of fluid exposed to the steep transient concentration gradients present initially after formation of the interface between the pure fluids is limited.

The properties of a microfluidic free interface created in accordance with embodiments of the present invention may be contrasted with a non-free microfluidic interface, as illustrated in FIGS. 33A and 33B. Specifically, FIG. 33A shows a simplified cross-sectional view of a microfluidic free interface in accordance with an embodiment of the present invention. Microfluidic free interface 7500 of FIG. 33A is formed between first fluid A and second fluid B present within channel 7502. The free microfluidic interface 7500 is substantially linear, with the result that the steep concentration gradient arising between fluids A and B is highly localized within the channel.

As described above, the dimensions of channel 7502 are extremely small, with the result that non-slip layers immediately adjacent to the walls of the channel in fact occupy most of the volume of the channel. As a result, viscosity forces are much greater than buoyant forces, and mixing between fluids A and B along interface 7500 occurs almost entirely as a result of diffusion, with little or no convective mixing. Conditions associated with the microfluidic free interface of embodiments of the present invention can be expressed in terms of the Grashof number (Gr) per Equation (2) below, an expression of the relative magnitude of buoyant and viscous forces: Gr∝B/V,  (2) where:

-   -   Gr=Grashof number;     -   B=buoyancy force; and     -   V=viscous force.         Microfluidic free interfaces in accordance with embodiments of         the present invention would be expected to exhibit a Grashof         number of 1 or less. The Grashof number expected with two fluids         having the same density is zero, and thus Grashof, numbers very         close to zero would be expected to be attained.

The embodiment of a microfluidic free interface illustrated above in FIG. 33A may be contrasted with the conventional non-microfluidic free interface shown in FIG. 33B. Specifically, first and second fluids A and B are separated by an interface 7504 that is not uniform or limited by the cross-sectional width of channel 7502. The steep concentration gradient occurring at the interface is not localized, but is instead present at various points along the length of the channel, exposing correspondingly large volumes of the fluids to the steep gradients. In addition, viscosity forces do not necessarily dominate over buoyancy forces, with the result that mixing of fluids A and B across interface 7504 can occur both as the result of diffusion and of convective flow. The Grashof number exhibited by such a conventional non-microfluidic interface would be expected to exceed 1.

TABLE A below provides a summary of the properties of a free microfluidic interface in accordance with an embodiment of the present invention, as contrasted with a non-microfluidic free interface.

TABLE A PROPERTY μFI non-μFI fluid forces dominated by a combination of viscous viscous force and buoyant forces Grashof No. low (i.e. <1) higher (i.e. >1) mechanism for mixing of predominantly a combination of diffusion fluids across interface diffusion and convective flow location of steepest localized non-localized concentration gradient between the fluids in their pure form approximate magnitude 100 nm-100 μm >100 μm of restrictive dimension at interface interface shape planar non-planar

1. Creation of Microfluidic Free Interface

A microfluidic free interface in accordance with embodiments of the present invention may be created in a variety of ways. One approach is to utilize the microfabricated elastomer structures previously described. Specifically, in certain embodiments the elastomeric material from which microfluidic structures are formed is relatively permeable to certain gases. This gas permeability property may be exploited utilizing the technique of pressurized out-gas priming (POP) to form well-defined, reproducible fluidic interfaces.

FIG. 34A shows a plan view of a flow channel 9600 of a microfluidic device in accordance with an embodiment of the present invention. Flow channel 9600 is separated into two halves by actuated valve 9602. Prior to the introduction of material, flow channel 9600 contains a gas 9604.

FIG. 34B shows the introduction of a first solution 9606 to first flow channel portion 9600 a under pressure, and the introduction of a second solution 9608 to second flow channel portion 9600 b under pressure. Because of the gas permeability of the surrounding elastomer material 9607, gas 9604 is displaced by the incoming solutions 9608 and 9610 and out-gasses through elastomer 9607.

As shown in FIG. 34C, the pressurized out-gas priming of flow channel portions 9600 a and 9600 b allows uniform filling of these dead-ended flow channel portions without air bubbles. Upon deactuation of valve 9602 as shown in FIG. 34D, microfluidic free interface 9612 is defined, allowing for formation of a diffusion gradient between the fluids.

While the specific embodiment just described exploits the permeability of the bulk material to dead end fill two or more chambers or channels separated by a closed valve, and creates a microfluidic free interface between the static fluids by the subsequent opening of this valve, other mechanisms for realizing a microfluidic free interface are possible.

For example, FIG. 47 shows one potential alternative method for establishing a microfluidic free interface diffusion in accordance with the present invention. Microfluidic channel 8100 carries fluid A experiencing a convective flow in the direction indicated by the arrow, such that static non-slip layers 8102 are created along the walls of flow channel 8100. Branch channel 8104 and dead-ended chamber 8106 contain static fluid B. Because material surrounding the dead-ended channel and chamber provide a back pressure, fluid B remains static and microfluidic free interface 8110 is created at mouth 8112 of branch channel 8108 between flowing fluid A and static fluid B. As described below, diffusion of fluid A or components thereof across the microfluidic free interface can be exploited to obtain useful results. While the embodiment shown in FIG. 47 includes a dead-ended branch channel and chamber, this is not required by the present invention, and the branch channel could connect with another portion of the device, as long as a sufficient counter pressure was maintained to prevent any net flow of fluid through the channel.

Another potential alternative method for establishing a microfluidic free interface diffusion assay is the use of break-through valves and chambers. A break-through valve is not a true closing valve, but rather a structure that uses the surface tension of the working fluid to stop the advance of the fluid. Since these valves depend on the surface tension of the fluid they can only work while a free surface exists at the valve; not when the fluid continuously fills both sides and the interior of the valve structure.

A non-exclusive list of ways to achieve such a valve include but are not limited to patches of hydrophobic material, hydrophobic treatment of certain areas, geometric constrictions (both in height and width) of a channel, geometric expansions (both in height and in width of a channel), changes in surface roughness on walls of a channel, and applied electric potentials on the walls.

These “break-through” valves may be designed to withstand a fixed and well defined pressure before they “break through” and allow fluid to pass nearly unimpeded. The pressure in the channel can be controlled and hence the fluid can be caused to advance when desired. Different methods of controlling this pressure include but are not limited to externally applied pressure at an input or output port, pressure derived from centrifugal force (i.e. by spinning the device), pressure derived from linear acceleration (i.e. applying an acceleration to the device with a component parallel to the channel), electrokinetic pressure, internally generated pressure from bubble formation (by chemical reaction or by hydrolysis), pressure derived from mechanical pumping, or osmotic pressure.

“Break-through” valves may be used to create a microfluidic free interface as shown and described in connection with FIGS. 35A-E. FIG. 35A shows a simplified plan view of a device for creating a microfluidic free interface utilizing break through valves. First chamber 9100 is in fluid communication with second chamber 9102 through branches 9104 a and 9104 b respectively, of T-shaped channel 9104.

First break through valve 9106 is located at outlet 9108 of first chamber 9100. Second break through valve 9110 is located in branch 9104 b upstream of inlet 9105 of second chamber 9102. Third break through valve 9112 is located at outlet 9114 of second chamber 9102. Breakthrough valves 9106, 9110, and 9112 may be formed from hydrophobic patches, a constriction in the width of the flow channel, or some other way as described generally above. In FIGS. 35A-E, an open break through valve is depicted as an unshaded circle, and a closed break through valve is depicted as a shaded circle.

In the initial stage shown in FIG. 35B, first chamber 9100 is charged with first fluid 9116 introduced through stem 9104 c and branch 9104 a of T-shaped channel 9104 and chamber inlet 9107 at a pressure below the break through pressure of any of the valves 9106, 9110, and 9112. In the second stage shown in FIG. 35C, second chamber 9102 is charged with a buffer or other intermediate fluid 9118 introduced through stem 9104 c of T-shaped channel 9104 and inlet 9105 at a pressure below the break through pressure of valve 9106 but greater than the break through pressures of valves 9110 and 9112.

In the third stage shown in FIG. 35D, intermediate fluid 9118 is replaced in second chamber 9102 with second fluid 9120 introduced through stem 9104 c of T-shaped channel 9104 and inlet 9105 at a pressure below the break through pressure of valve 9106 but greater than the break through pressures of valves 9110 and 9112. In the final stage depicted in FIG. 35E, second fluid 9120 has replaced the intermediate fluid, leaving the first and second fluids 9116 and 9120 in separate chambers but fluidically connected through T-junction 9104, creating a microfluidic free interface 9122.

The use of break through valves to create a microfluidic free interface in accordance with embodiments of the present invention is not limited to the specific example given above. For example, in alternative embodiments the step of flushing with a buffer or intermediate solution is not required, and the first solution could be removed by flushing directly with the second solution, with potential unwanted by-products of mixing removed by the initial flow of the second solution through the channels and chambers.

While the embodiments just described create the microfluidic free interface in a closed microfluidic device, this is not required by embodiments in accordance with the present invention. For example, an alternative embodiment in accordance with the present invention may utilize capillary forces to connect two reservoirs of fluid. In one approach, the open wells of a micro-titer plate could be connected by a segment of a glass capillary. The first solution would be dispensed into one well such that it fills the well and is in contact with the glass capillary. Capillary forces cause the first solution to enter and flow to the end of the capillary. Once at the end, the fluid motion ceases. Next, the second solution is added to the second well. This solution is in contact with the first solution at the capillary inlet and creates a microfluidic interface between the two wells at the end of the capillary.

The connecting path between the two wells need not be a glass capillary, and in alternative embodiments could instead comprise a strip of hydrophilic material, for example a strip of glass or a line of silica deposited by conventional CVD or PVD techniques. Alternatively, the connecting paths could be established by paths of less hydrophobic material between patterned regions of highly hydrophobic material. Moreover, there could be a plurality of such connections between the wells, or a plurality of interconnected chambers in various configurations. Such interconnections could be established by the user prior to use of the device, allowing for rapid and efficient variation in fluidic conditions.

Where as in the previous example the two reservoirs are not enclosed by a microfluidic device but are connected instead through a microfluidic path, an alternative embodiment could have reservoirs both enclosed and not enclosed. For example, sample could be loaded into a microfluidic device and pushed to the end of an exit capillary or orifice (by any of the pressure methods described above). Once at the end of the exit capillary, the capillary could be immersed in a reservoir of reagent. In this way, the microfluidic free interface is created between the external reservoir and the reservoir of reagent in the chip. This method could be used in parallel with many different output capillaries or orifices to screen a single sample against a plurality of different reagents using microfluidic free interface diffusion.

In the example just described, the reagent is delivered from one or many inlets to one or many different outlets “through” a microfluidic device. Alternatively, this reagent can be introduced through the same orifice that is to be used to create the microfluidic interface. The sample-containing solution could be aspirated into a capillary (either by applying suction, or by capillary forces, or by applying pressure to the solution) and then the capillary may be immersed in a reservoir of counter-reagent, creating a microfluidic interface between the end of the capillary and the reservoir. This could be done in a large array of capillaries for the parallel screening of many different reagents. Very small volumes of sample could be used since the capillaries can have a fixed length beyond which the sample will not advance. For crystallization applications (see below), the capillaries could be removed and mounted in an x-ray beam for diffraction studies, without requiring handling of the crystals.

2. Reproducible Control Over Equilibration Parameters

One advantage of the use of microfluidic free interface diffusion in accordance with embodiments of the present invention is the ability to create uniform and continuous concentration gradients that reproducibly sample a wide range of conditions. As the fluids on either side of the interface diffuse into one another, a gradient is established along the diffusion path, and a continuum of conditions is simultaneously sampled. Since there is a variation in the conditions, both in space and time, information regarding the location and time of positive results (i.e. crystal formation) may be used in further optimization.

In many applications it is desirable to create a gradient of a condition such as pH, concentration, or temperature. Such gradients may be used for screening applications, optimization of reaction conditions, kinetics studies, determination of binding affinities, dissociation constants, enzyme-rate profiling, separation of macromolecules, and many other applications. Due principally to the suppression of convective flow, diffusion across a microfluidic free interface in accordance with an embodiment of the present invention may be used to establish reliable and well-defined gradient.

The dimensional Einstein equation (3) may be employed to obtain a rough estimate of diffusion times across a microfluidic free interface.

$\begin{matrix} {{t = \frac{x^{2}}{4D}};} & (3) \end{matrix}$ where:

-   -   t=diffusion time;     -   x=longest diffusion length; and     -   D=diffusion coefficient         Generally, as shown in Equation (4) below, the diffusion         coefficient varies inversely with the radius of gyration, and         therefore as one over the cube root of molecular weight.

$\begin{matrix} {{D \propto \frac{1}{r} \propto \frac{1}{m^{1/3}}};} & (4) \end{matrix}$ where:

-   -   D=diffusion coefficient;     -   r=radius of gyration; and     -   m=molecular weight         In reviewing equation (4), it is important to recognize that         correlation between the radius of gyration (r) and the molecular         weight (m) is only an approximation. Because of the dominance of         viscous forces over inertial forces, the diffusion coefficient         is in fact independent of molecular weight and is instead         dependent upon the size and hence drag experienced by the         diffusing particle.

As compared with the rough 1.5 hr equilibration time for a dye, an approximate equilibration time for a protein of 20 KDa over the same distance is estimated to be approximately 45 hours. The equilibration time for a small salt of a molecular weight of 100 Da over the same distance is about 45 minutes.

FIGS. 36A-D are simplified schematic diagrams plotting concentration versus distance for a solution A and a solution B in contact along a free interface. FIGS. 36A-D show that over time, a continuous and broad range of concentration profiles of the two solutions is ultimately created.

Moreover, varying the equilibration rate by changing the geometry over time of connecting channels may be used on a single device to explore the effect of equilibration dynamics. FIGS. 37A-D show an embodiment in which a gradient of concentrations, initially established by the partial diffusive equilibration of two solutions from a micro-free interface, can be captured by actuation of containment valves.

FIG. 37A shows flow channel 10000 that is overlapped at intervals by a forked control channel 10002 to define a plurality of chambers (A-G) positioned on either side of a separately-actuated interface valve 10004. FIG. 37B plots solvent concentration at an initial time, when interface valve 10004 is actuated and a first half 10000 a of the flow channel has been mixed with a first solution, and a second half 10000 b of the flow channel has been primed with a second solution. FIG. 37C plots solvent concentration at a subsequent time T₁, when control channel 10002 is actuated to define the seven chambers (A-G), which capture the concentration gradient at that particular point in time. FIG. 37D plots relative concentration of the chambers (A-G) at time T₁.

In the embodiment shown in FIG. 37A, actuation of the forked control channel simultaneously creates the plurality of chambers A-G. However, this is not required, and in alternative embodiments of the present invention multiple control channels could be utilized to allow independent creation of chambers A-G at different time intervals, thereby allow additional diffusion to occur after an initial set of chambers are created immediately adjacent to the free interface.

The relative concentrations resulting from diffusion across a fluidic interface is determined not only by thermodynamic conditions explored during the equilibration, but also by the rate at which equilibration takes place. It is therefore potentially valuable to control the dynamics of equilibration.

In conventional macroscopic diffusion methods, only coarse control over the dynamics of equilibration may be available through manipulation of initial conditions. For macroscopic free interface diffusion, once diffusion begins, the experimenter has no control over the subsequent equilibration rate. For hanging drop experiments, the equilibration rate may be changed by modifying the size of the initial drop, the total size of the reservoir, or the temperature of incubation. In microbatch experiments, the rate at which the sample is concentrated may be varied by manipulating the size of the drop, and the identity and amount of the surrounding oil. Since the equilibration rates depend in a complicated manner on these parameters, the dynamics of equilibration may only be changed in a coarse manner. Moreover, once the experiment has begun, no further control over the equilibration dynamics is available.

By contrast, in a fluidic free interface experiment in accordance with an embodiment of the present invention, the parameters of diffusive equilibration rate may also be controlled by manipulating dimensions of chambers and connecting channels of a microfluidic structure. For example, in a microfluidic structure comprising reservoirs in fluid communication through a constricted channel, where no appreciable gradient exists in the reservoirs due to high concentrations or replenishment of material, to good approximation the time required for equilibration varies linearly with the required diffusion length. The equilibration rate also depends on the cross-sectional area of the connecting channels. The required time for equilibration may therefore be controlled by changing both the length, and the cross-sectional area of the connecting channels.

For example, FIG. 40A shows a plan view of a simple embodiment of a microfluidic structure in accordance with the present invention. Microfluidic structure 9701 comprises reservoirs 9700 and 9702 containing first fluid A and second fluid B, respectively. Reservoirs 9700 and 9702 are connected by channel 9704. Valve 9706 is positioned on the connecting channel between reservoirs 9700 and 9702.

Connecting channel 9704 has a much smaller cross-sectional area than either of the reservoirs. For example, in particular embodiments of microfluidic structures in accordance with the present invention, the ratio of reservoir/channel cross-sectional area and thus the ratio of maximum ratio of cross-sectional area separating the two fluids, may fall between 500 and 25,000. The minimum of this range describes a 50×50×50 μm chamber connected to a 50×10 μm channel, and the maximum of this range describes a 500×500×500 μm chamber connected to a 10×1 μm channel.

Initially, reservoirs 9700 and 9702 are filled with respective fluids, and valve 9706 is closed. Upon opening valve 9706, a microfluidic free interface in accordance with an embodiment of the present invention is created, and fluids A and B diffuse across this interface through the channel into the respective reservoirs. Moreover, where the amount of diffusing material present in one reservoir is large and the capacity of the other reservoir to receive material without undergoing a significant concentration change is also large, the concentrations of material in the reservoirs will not change appreciably over time, and a steady state of diffusion will be established.

Diffusion of fluids in the simple microfluidic structure shown in FIG. 40A may be described by relatively simple equations. For example, the net flux of a chemical species from one chamber to the other may be simply described by equation (5):

$\begin{matrix} {{J = {D*A*\frac{\Delta\; C}{L}}};} & (5) \end{matrix}$ where:

-   -   J=net flux of chemical species     -   D=diffusion constant of the chemical species;     -   A=cross-sectional area of the connecting channel;     -   ΔC=concentration difference between the two channels; and     -   L=length of the connection channel.

Following integration and extensive manipulation of the terms of equation (5), the characteristic time τ for the equilibration of the two chambers, where one volume V₁ is originally at concentration C and the other volume V₂ is originally at concentration 0, can therefore be taken to be as shown in Equation (6) below:

$\begin{matrix} {{\tau = {\frac{1}{{V_{1}/V_{2}} + 1}*\frac{1}{D}*\frac{L}{A/V_{1}}}};} & (6) \end{matrix}$ where:

-   -   τ=equilibration time;     -   V₁=volume of chamber initially containing the chemical species;     -   V₂=volume of chamber into which the chemical species is         diffusing;     -   D=diffusion constant of the chemical species;     -   A=cross-sectional area of the connecting channel; and     -   L=length of the connection channel.

Therefore, for a given initial concentration of a chemical species in a chamber of a defined volume, the characteristic equilibration time depends in a linear manner from the diffusive length L and the ratio of the cross-sectional area to the volume (hereafter referred to simply as the “area”), with the understanding that the term “area” refers to the area normalized by the volume of the relevant chamber. Where two chambers are connected by a constricted channel, as in the structure of FIG. 40A, the concentration drop from one channel to the other occurs primarily along the connecting channel and there is no appreciable gradient present in the chamber. This is shown in FIG. 40B, which is a simplified plot of concentration versus distance for the structure of FIG. 40A.

The behavior of diffusion between the chambers of the microfluidic structure of FIG. 40A can be modeled, for example, utilizing the PDE toolbox of the MATLAB® software program sold by The MathWorks Inc. of Natick, Mass. FIGS. 41 and 42 accordingly show the results of simulating diffusion of sodium chloride from a 300 um×300 um×100 um chamber to another chamber of equal dimensions, through a 300 um long channel with a cross-sectional area of 1000 um. The initial concentrations of the chambers are 1 M and 0 M, respectively.

FIG. 41 plots the time required for the concentration in one of the reservoirs to reach 0.6 of the final equilibration concentration, versus channel length. FIG. 41 shows the linear relationship between diffusion time and channel length for this simple microfluidic system.

FIG. 42 plots the inverse of the time required for the concentration in one of the reservoirs to reach 0.6 of the final equilibration concentration (T_(0.6)), versus the area of the fluidic interface created upon opening of the valve. FIG. 42 shows the linear relationship between these parameters. The simple relationship between the equilibration time constant and the parameters of channel length and 1/channel area allows for a reliable and intuitive method for controlling the rate of diffusive mixing across a microfluidic free interface in accordance with an embodiment of the present invention.

This relationship further allows for one reagent to be diffusively mixed with a plurality of others at different rates that may be controlled by the connecting channel geometry. For example, FIG. 38A shows three sets of pairs of compound chambers 9800, 9802, and 9804, each pair connected by microchannels 9806 of a different length Δx. FIG. 38B plots equilibration time versus equilibration distance. FIG. 38B shows that the required time for equilibration of the chambers of FIG. 38A varies as the length of the connecting channels.

FIG. 39 shows four compound chambers 9900, 9902, 9904, and 9906, each having different arrangements of connecting microchannel(s) 9908. Microchannels 9908 have the same length, but differ in cross-sectional area and/or number of connecting channels. The rate of equilibration may thus be increased/decreased by decreasing/increasing the cross-sectional area, for example by decreasing/increasing the number of connecting channels or the dimensions of those channels.

Another desirable aspect of microfluidic free interface diffusion studies in accordance with embodiments of the present invention is the ability to reproducibly explore a wide range of phase space. For example, it may be difficult to determine, a priori, which thermodynamic conditions will be favorable for a particular application (i.e. nucleation/growth of protein crystals), and therefore it is desirable that a screening method sample as much of phase-space (as many conditions) as possible. This can be accomplished by conducting a plurality of assays, and also through the phase space sampled during the evolution of each assay in time.

FIG. 43 shows the results of simulating the counter-diffusion of lysosyme and sodium chloride utilizing the microfluidic structure shown in FIG. 40A, with different relative volumes of the two reservoirs, and with initial concentrations normalized to 1. FIG. 43 presents a phase diagram depicting the phase space between fluids A and B, and the path in phase space traversed in the reservoirs as the fluids diffuse across the microfluidic free interface created by the opening of the valve in FIG. 40A. FIG. 43 shows that the phase space sampled depends upon the initial relative volumes of the fluids contained in the two reservoirs. By utilizing arrays of chamber with different sample volumes, and then identifying instances where diffusion across the channel yielded desirable results (i.e. crystal formation), promising starting points for additional experimentation can be determined.

As described above, varying the length or cross-sectional area of a channel connecting two reservoirs changes the rate at which the species are mixed. However, so long as the channel volume remains small compared as compared with the total reaction volume, there is little or no effect on the evolution of concentration in the chambers through phase space. The kinetics of the mixing are therefore decoupled from the phase-space evolution of the reaction, allowing the exercise independent control over the kinetic and thermodynamic behavior of the diffusion.

For example, it is often desirable in crystallography to slow down the equilibration so as to allow for the growth of fewer and higher quality crystals. In conventional techniques this is often attempted by adding new chemical constituents such as glycerol, or by using microbatch methods. However, this addition of constituents is not well characterized, is not always effective, and may inhibits the formation of crystals. Microbatch methods also may pose the disadvantage of lacking a driving force to promote continued crystal growth as protein in the solution surrounding the crystal is depleted. Through the use of diffusion across a microfluidic free interface in accordance with an embodiment of the present invention, crystal formation may be slowed by a well-defined amount without altering the phase-space evolution, simply by varying the width or cross-sectional area of the connecting channel.

The ability to control the rate at which equilibration proceeds has further consequences in cases were one wishes to increase the total volume of a reaction while conserving both the thermodynamics and the microfluidic free interface diffusion mixing. One such case arises again in the context of protein crystallography, in which an initial, small volume crystallization assay results in crystals of insufficient size for diffraction studies. In such a case, it is desirable to increase the reaction volume and thereby provide more protein available for crystal growth, while at the same time maintaining the same diffusive mixing and path through phase space. By increasing the chamber volumes proportionally and decreasing the area of the channel, the area of the interface relative to the total assay volume is reduced, and a larger volume would pass through the same phase space as in the original small volume conditions.

While the above description has focused upon diffusion of a single species, gradients of two or more of species which do not interact with each other may be created simultaneously and superimposed to create an array of concentration conditions. FIG. 52 shows a plan view of one example of a microfluidic structure for creating such superimposed gradients. Flat, shallow chamber 8600 constricted in the vertical direction is connected at its periphery to reservoirs 8602 and 8604 having fixed concentrations of chemical species A and B, respectively. Sink 8606 in the form of a reservoir is maintained at a substantially lower concentration of species A and B. After the initial transient equilibration, stable and well-defined gradients 8608 and 8610 of species A and B respectively, are established in two dimensions.

As evident from inspection of FIG. 52, the precise shape and profile of the concentration gradient will vary according to a host of factors, including but not limited to the relative location and number of inlets to the chamber, which can also act as concentration sinks for the chemical species not contained therein (i.e. reservoir 8604 may act as a sink for chemical species A). However, the spatial concentration profiles of each chemical species within the chamber may readily be modeled using the MATLAB program previously described to describe a two-dimensional, well-defined, and continuous spatial gradient.

The specific embodiment illustrated in FIG. 52 offers the disadvantage of continuous diffusion of materials. Hence, where diffusion of products of reaction between the diffusing species is sought to be discerned, these products will themselves diffuse in the continuous gradient, thereby complicating analysis.

Accordingly, FIG. 53 shows a simplified plan view of an alternative embodiment of a microfluidic structure for accomplishing diffusion in two dimensions. Grid 8700 of intersecting orthogonal channels 8702 establishes a spatial concentration gradient. Reservoirs 8704 and 8708 of fixed concentrations of chemical species A and B are positioned on adjacent edges of grid 8700. Opposite to these reservoirs on the grid are two sinks 8703 of lower concentration of the chemical present in the opposing reservoirs.

Surrounding each channel junction 8710 are two pairs of valves 8712 and 8714 which control diffusion through the grid in the vertical and horizontal directions, respectively. Initially, only valve pairs 8714 are opened to create a well-defined diffusion gradient of the first chemical in the horizontal direction. Next, valve pairs 8714 are closed and valve pairs 8712 opened to create a well-defined diffusion gradient of the second chemical in the vertical direction. Isolated by adjacent horizontal valves, the gradient of the first chemical species remains present in regions between the junctions.

Once the second (vertical) gradient is established, the two gradients can be combined and by opening all the valve pairs for a short time to allow partial diffusive equilibration. After the period of diffusion has passed, all the valve pairs are closed to contain the superimposed gradient. Alternatively, valve pairs 8712 and 8714 can be closed to halt diffusion in the vertical direction, with every second horizontal valve opened to create separate isolated chambers.

3. Control Over Environmental Factors

The above discussion has focused upon altering the chemical environment through diffusive mixing across a microfluidic free interface between two fluids. Other factors, however, may also influence interaction between the fluids on either side of the microfluidic free interface. Such additional factors include, but are not limited to, temperature, pressure, and the concentration of materials in the fluids.

In specific embodiments in accordance with the present invention, control over temperature during diffusive mixing may be accomplished utilizing a composite elastomer/silicon structure previously described. Specifically, a Peltier temperature control structure may be fabricated in an underlying silicon substrate, with the elastomer aligned to the silicon such that a crystallization chamber is proximate to the Peltier device. Application of voltage of an appropriate polarity and magnitude to the Peltier device may control the temperature of fluids within a chamber or channel.

Alternatively, as described by Wu et al. in “MEMS Flow Sensors for Nano-fluidic Applications”, Sensors and Actuators A 89 152-158 (2001), microfabricated chambers or channels could be heated and cooled through the selective application of current to a micromachined resistor structure resulting in ohmic heating. Moreover, the temperature of the fluid could be detected by monitoring the resistance of the heater over time. The Wu et al. paper is hereby incorporated by reference for all purposes.

It may also be useful to establish a temperature gradient across a microfabricated free interface in accordance with the present invention. Such a temperature gradient would subject the diffusing fluids to a broad spectrum of temperatures during mixing, allowing for extremely precise determination of optimum temperatures.

With regard to controlling pressure during diffusive mixing, embodiments of the present invention which meter fluids by volume exclusion are particularly advantageous. Specifically, once the chamber has been charged with appropriate fluid volumes, a chamber inlet valve may be maintained shut while the membrane overlying the chamber is actuated, thereby causing pressure to increase in the chamber. Structures in accordance with the present invention employing techniques other than volume exclusion could exert pressure control by including flow channels and associated membranes adjacent to the diffusion channel or chamber that are specifically relegated to controlling pressure.

III. Applications for Microfluidic Free Interface Diffusion

Microfluidic free interfaces created in accordance with embodiments of the present invention may be useful in a variety of applications. As specifically described in detail below, the exclusively diffusive mixing that occurs at such a microfluidic free interface may provide conditions advantageous in the formation of crystals, as well as in the performance of certain assays.

In general, microfluidics enables the handling of fluids on the sub-nanoliter scale. Consequently, there is no need to use large containment chambers, and hence, assays may be performed on the nanoliter, or subnanoliter scale. The utilization of extremely small volumes allows for thousands of assays to be performed while consuming the same sample volume required for one macroscopic free-interface diffusion experiment. This reduces costly and time-consuming amplification and purification steps, and makes possible the screening of proteins that are not easily expressed, and hence must be purified from a bulk sample.

Microfluidics further offers savings in preparation times, as hundreds, or even thousands of assays may be performed simultaneously. The use of scaleable metering techniques as previously described, allow for parallel experimentation to be conducted without increased complexity in control mechanisms.

Finally, other researchers have proposed the diffusion of materials across an interface between two flowing liquids. Kamholz et al., “Quantitative Analysis of Molecular Interaction in a Microfluidic Channel”, Anal. Chem. Vol. 71, No. 20 (1999). However, such an approach requires relatively large volumes of sample owing to the continuous flow of materials, and also involves the exposure of large volumes of the flowing liquid to the relatively steep concentration gradient present at the fluidic interface. Embodiments in accordance with the present invention utilizing diffusion across a microfluidic free interface avoid these and other problems associated with diffusion between flowing materials.

1. Protein Crystallography

The use of microfluidic free interface diffusion techniques in protein crystallography has previously been discussed in detail in U.S. nonprovisional patent application Ser. No. 09/887,997 filed Jun. 22, 2001; and U.S. nonprovisional patent application Ser. No. 10/117,978 filed Apr. 5, 2002. These applications are hereby incorporated by reference for all purposes.

As employed in the following discussion, the term “crystallizing agent” describes a substance that is introduced to a solution of target material to lessen solubility of the target material and thereby induce crystal formation. Crystallizing agents typically include countersolvents in which the target exhibits reduced solubility, but may also describe materials affecting solution pH or materials such as polyethylene glycol that effectively reduce the volume of solvent available to the target material. The term “countersolvent” is used interchangeably with “crystallizing agent”.

Briefly, crystallization is an important technique to the biological and chemical arts. A high-quality crystal of a target compound can be analyzed by x-ray diffraction techniques to produce an accurate three-dimensional structure of the target. This three-dimensional structure information can then be utilized to predict functionality and behavior of the target.

In theory, the crystallization process is simple. A target compound in pure form is dissolved in or exchanged into some type of solvent. For proteins, this solvent is typically, although not always, water. For small molecule chemicals, the solvent can be water- or organic-based. The chemical environment of the dissolved target material is then altered such that the target is less soluble and reverts to the solid phase in crystalline form. This change in chemical environment typically accomplished by introducing a crystallizing agent that makes the target material is less soluble, although changes in temperature and pressure can also influence solubility of the target material.

In practice however, forming a high quality crystal is generally difficult and sometimes impossible, requiring much trial and error and patience on the part of the researcher. Specifically, the highly complex structure of even simple biological compounds means that they are not amenable to forming a highly ordered crystalline structure. Therefore, a researcher must be patient and methodical, experimenting with a large number of conditions for crystallization, altering parameters such as sample concentration, solvent type, countersolvent type, temperature, and duration in order to obtain a high quality crystal, if in fact a crystal can be obtained at all.

Systems and methods in accordance with embodiments of the present invention are particularly suited to crystallizing larger biological macromolecules or complexes thereof, such as proteins, nucleic acids, viruses, and protein/ligand assemblies. However, crystallization in accordance with the present invention is not limited to any particular type of target material.

In accordance with embodiments of the present invention, a crystallization technique called gated micro free interface diffusion (Gated μ-FID), has been developed. Referring back to FIGS. 34A-D, one generic embodiment of protein crystallization utilizing microfluidic free interface diffusion introduces a sample solution 9606 into first flow channel portion 9600 a under pressure, and introduces a solution containing a countersolvent or crystallizing agent into second flow channel portion 9600 b under pressure. Because of the gas permeability of the surrounding elastomer material 9607, gas 9604 is displaced by the incoming solutions 9608 and 9610 and out-gasses through elastomer 9607.

As shown in FIG. 34C, the pressurized out-gas priming of flow channel portions 9600 a and 9600 b allows uniform filling of these dead-ended flow channel portions without air bubbles. Upon deactuation of valve 9602 as shown in FIG. 34D, a microfluidic free interface 9612 is defined, thereby allowing for formation of a diffusion gradient between the sample solution and the solution containing the crystallizing agent. As a result of diffusive mixing between the sample and the crystallizing agent, the solution environment is gradually changed, resulting in the formation of protein crystals in the channel.

The formation of protein crystals utilizing gated μ-FID retains the efficient sampling of phase space achieved by macroscopic free interface diffusion techniques, with a number of added advantages, including the parsimonious use of sample solutions, ease of set-up, creation of well defined fluidic interfaces, control over equilibration dynamics, and the ability to conduct high-throughput parallel experimentation.

Another possible advantage of the formation of protein crystals utilizing gated μ-FID is the formation of high quality crystals, as illustrated in connection with FIGS. 44A and 44B. FIG. 44A shows a simplified schematic view of a protein crystal being formed utilizing a conventional macroscopic free interface diffusion technique. Specifically, nascent protein crystal 9200 is exposed to sample from solution 9202 that is experiencing a net conductive flow of sample as a result of the action of buoyant forces. As a result of the directionality of this conductive flow, the growth of protein crystal 9200 is also directional. However, as described in Nerad et al., “Ground-Based Experiments on the Minimization of Convection During the Growth of Crystals from Solution”, Journal of Crystal Growth 75, 591-608 (1986), asymmetrical growth of a protein crystal can give rise to unwanted strain in the lattice of the growing crystal, promoting dislocations and/or the incorporation of impurities within the lattice, and otherwise adversely affecting the crystal quality.

By contrast, FIG. 44B shows a simplified schematic view of a protein crystal being formed utilizing diffusion across a microfluidic free interface in accordance with an embodiment of the present invention. Nascent protein crystal 9204 is exposed to sample solution 9206 that is diffusing within the crystallizing agent. This diffusion is nondirectional, and the growth of protein crystal 9204 is also correspondingly nondirectional. Accordingly, the growing crystal avoids strain on the lattice and the attendant incorporation of impurities and dislocations experienced by the growing crystal shown in FIG. 44A. Accordingly, the quality of the crystal in FIG. 44B is of high quality.

Typical targets for crystallization are diverse. A target for crystallization may include but is not limited to: 1) biological macromolecules (cytosolic proteins, extracellular proteins, membrane proteins, DNA, RNA, and complex combinations thereof), 2) pre- and post-translationally modified biological molecules (including but not limited to, phosphorylated, sulfolated, glycosylated, ubiquitinated, etc. proteins, as well as halogenated, abasic, alkylated, etc. nucleic acids); 3) deliberately derivatized macromolecules, such as heavy-atom labeled DNAs, RNAs, and proteins (and complexes thereof), selenomethionine-labeled proteins and nucleic acids (and complexes thereof), halogenated DNAs, RNAs, and proteins (and complexes thereof), 4) whole viruses or large cellular particles (such as the ribosome, replisome, spliceosome, tubulin filaments, actin filaments, chromosomes, etc.), 5) small-molecule compounds such as drugs, lead compounds, ligands, salts, and organic or metallo-organic compounds, and 6) small-molecule/biological macromolecule complexes (e.g., drug/protein complexes, enzyme/substrate complexes, enzyme/product complexes, enzyme/regulator complexes, enzyme/inhibitor complexes, and combinations thereof). Such targets are the focus of study for a wide range of scientific disciplines encompassing biology, biochemistry, material sciences, pharmaceutics, chemistry, and physics.

During crystallization screening, a large number of chemical compounds may be employed. These compounds include salts, small and large molecular weight organic compounds, buffers, ligands, small-molecule agents, detergents, peptides, crosslinking agents, and derivatizing agents. Together, these chemicals can be used to vary the ionic strength, pH, solute concentration, and target concentration in the drop, and can even be used to modify the target. The desired concentration of these chemicals to achieve crystallization is variable, and can range from nanomolar to molar concentrations. A typical crystallization mix contains set of fixed, but empirically-determined, types and concentrations of ‘precipitants’, buffers, salts, and other chemical additives (e.g., metal ions, salts, small molecular chemical additives, cryo protectants, etc.). Water is a key solvent in many crystallization trials of biological targets, as many of these molecules may require hydration to stay active and folded.

‘Precipitating’ agents act to push targets from a soluble to insoluble state, and may work by volume exclusion, changing the dielectric constant of the solvent, charge shielding, and molecular crowding. Precipitating agents compatible with the PDMS material of certain embodiments of the chip include, but are not limited to, salts, high molecular weight polymers, polar solvents, aqueous solutions, high molecular weight alcohols, and divalent metals.

Precipitating compounds, which include large and small molecular weight organics, as well as certain salts are used from under 1% to upwards of 40% concentration, or from less than 0.5M to greater than 4M concentration. Water itself can act in a precipitating manner for samples that require a certain level of ionic strength to stay soluble. Many precipitants may also be mixed with one another to increase the chemical diversity of the crystallization screen. The microfluidics devices described in this document are readily compatible with a broad range of such compounds. Moreover, many precipitating agents (such as long- and short-chain organics) are quite viscous at high concentrations, presenting a problem for most fluid handling devices, such as pipettes or robotic systems. The pump and valve action of microfluidics devices in accordance with embodiments of the present invention enable handling of viscous agents.

An investigation of solvent/precipitating agent compatibility with particular elastomer materials may be conducted to identify optimum crystallizing agents, which may be employed develop crystallization screening reactions tailored for the chip that are more effective than standard screens.

Solution pH can be varied by the inclusion of buffering agents; typical pH ranges for biological materials lie anywhere between values of 3.5-10.5 and the concentration of buffer, generally lies between 0.01 and 0.25 M. The microfluidics devices described in this document are readily compatible with a broad range of pH values, particularly those suited to biological targets.

A nonexclusive list of possible buffers is as follows: Na/K-Acetate; HEPES; Na-Cacodylate; Na/K-Citrate; Na/K-Succinate; Na/K-Phosphate; TRIS; TRIS-Maleate; Imidazole-Maleate; BisTrisPropane; CAPSO, CHAPS, MES, and imidizole.

Additives are small molecules that affect the solubility and/or activity behavior of the target. Such compounds can speed crystallization screening or produce alternate crystal forms of the target. Additives can take nearly any conceivable form of chemical, but are typically mono and polyvalent salts (inorganic or organic), enzyme ligands (substrates, products, allosteric effectors), chemical crosslinking agents, detergents and/or lipids, heavy metals, organo-metallic compounds, trace amounts of precipitating agents, and small and large molecular weight organics.

Many of these chemicals can be obtained in predefined screening kits from a variety of vendors, including but not limited to Hampton Research of Laguna Niguel, Calif., Emerald Biostructures of Bainbridge Island, Wash., and Jena BioScience of Jena, Germany, that allow the researcher to perform both ‘sparse matrix’ and ‘grid’ screening experiments. Sparse matrix screens attempt to randomly sample as much of precipitant, buffer, and additive chemical space as possible with as few conditions as possible. Grid screens typically consist of systematic variations of two or three parameters against one another (e.g., precipitant concentration vs. pH). Both types of screens have been employed with success in crystallization trials, and the majority of chemicals and chemical combinations used in these screens are compatible with the chip design and matrices in accordance with embodiments of the present invention.

Moreover, current and future designs of microfluidic devices may enable flexibly combinatorial screening of an array of different chemicals against a particular target or set of targets, a process that is difficult with either robotic or hand screening. This latter aspect is particularly important for optimizing initial successes generated by first-pass screens.

In addition to chemical variability, a host of other parameters can be varied during crystallization screening. Such parameters include but are not limited to: 1) volume of crystallization trial, 2) ratio of target solution to crystallization solution, 3) target concentration, 4) cocrystallization of the target with a secondary small or macromolecule, 5) hydration, 6) incubation time, 7) temperature, 8) pressure, 9) contact surfaces, 10) modifications to target molecules, and 11) gravity.

Volumes of crystallization trials can be of any conceivable value, from the picoliter to milliliter range. Typical values may include but are not limited to: 0.1, 0.2, 0.25, 0.4, 0.5, 0.75, 1, 2, 4, 5, 10, 15, 20, 25, 30, 35, 40, 45, 50, 60, 70, 75, 80, 90, 100, 125, 150, 175, 200, 225, 250, 275, 300, 350, 400, 450, 500, 550, 600, 700, 750, 800, 900, 1000, 1100, 1200, 1250, 1300, 1400, 1500, 1600, 1700, 1800, 1900, 2000, 2250, 2500, 3000, 4000, 5000, 6000, 7000, 7500, 8000, 9000, and 10000 nL. The microfluidics devices previously described can access these values.

In particular, access to the low volume range for crystallization trials (<100 nL) is a distinct advantage of embodiments of the microfluidics chips in accordance with embodiments of the present invention, as such small-volume crystallization chambers can be readily designed and fabricated, minimizing the need the need for large quantities of precious target molecules. The low consumption of target material of embodiments in accordance with the present invention is particularly useful in attempting to crystallize scarce biological samples such as membrane proteins, protein/protein and protein/nucleic acid complexes, and small-molecule drug screening of lead libraries for binding to targets of interest.

The ratios of a target solution to crystallization mix can also constitute an important variable in crystallization screening and optimization. These rations can be of any conceivable value, but are typically in the range of 1:100 to 100:1 target:crystallization-solution. Typical target:crystallization-solution or crystallization-solution:target ratios may include but are not limited to: 1:100, 1:90, 1:80, 1:70, 1:60, 1:50, 1:40, 1:30, 1:25, 1:20, 1:15, 1:10, 1:9, 1:8, 1:7.5, 1:7, 1:6, 1:5, 1:4, 1:3, 1:2.5, 1:2, 1:1, 2:3, 3:4, 3:5, 4:5, 5:6, 5:7, 5:9, 6:7, 7:8, 8:9, and 9:10. As previously described, microfluidics devices in accordance with embodiments of the present invention can be designed to access multiple ratios simultaneously on a single chip.

Target concentration, like crystallization chemical concentration, can lie in a range of values and is an important variable in crystallization screening. Typical ranges of concentrations can be anywhere from <0.5 mg/ml to >100 mg/ml, with most commonly used values between 5-30 mg/ml. The microfluidics devices in accordance with embodiments of the present invention are readily compatible with this range of values.

Cocrystallization generally describes the crystallization of a target with a secondary factor that is a natural or nonnatural binding partner. Such secondary factors can be small, on the order of about 10-1000 Da, or may be large macromolecules. Cocrystallization molecules can include but are not limited to small-molecule enzyme ligands (substrates, products, allosteric effectors, etc.), small-molecule drug leads, single-stranded or double-stranded DNAs or RNAs, complement proteins (such as a partner or target protein or subunit), monoclonal antibodies, and fusion-proteins (e.g., maltose binding proteins, glutathione S-transferase, protein-G, or other tags that can aid expression, solubility, and target behavior). As many of these compounds are either biological or of a reasonable molecular weight, cocrystallization molecules can be routinely included with screens in the microfluidics chips. Indeed, because many of these reagents are expensive and/or of limited quantity, the small-volumes afforded by the microfluidics chips in accordance with embodiment of the present invention make them ideally suited for cocrystallization screening.

Hydration of targets can be an important consideration. In particular, water is by far the dominant solvent for biological targets and samples. The microfluidics devices described in this document are relatively hydrophobic, and are compatible with water-based solutions.

The length of time for crystallization experiments can range from minutes or hours to weeks or months. Most experiments on biological systems typically show results within 24 hours to 2 weeks. This regime of incubation time can be accommodated by the microfluidics devices in accordance with embodiments of the present invention.

The temperature of a crystallization experiment can have a great impact on success or failure rates. This is particularly true for biological samples, where temperatures of crystallization experiments can range from 0-42° C. Some of the most common crystallization temperatures are: 0, 1, 2, 4, 5, 8, 10, 12, 15, 18, 20, 22, 25, 30, 35, 37, and 42. Microfluidics devices in accordance with embodiments of the present invention can be stored at the temperatures listed, or alternatively may be placed into thermal contact with small temperature control structures such as resistive heaters or Peltier cooling structures.

In addition, the small footprint and rapid setup time of embodiments in accordance with the present invention allow faster equilibration to desired target temperatures and storage in smaller incubators at a range of temperatures. Moreover, as the microfluidics systems in accordance with embodiments of the present invention do not place the crystallization experiment in contact with the vapor phase, condensation of water from the vapor phase into the drop as temperatures change, a problem associated with conventional macroscopic vapor-diffusion techniques, is avoided. This feature represents an advance over many conventional manual or robotic systems, where either the system must be maintained at the desired temperature, or the experiment must remain at room temperature for a period before being transferred to a new temperature.

Variation in pressure is an as yet understudied crystallization parameter, in part because conventional vapor-diffusion and microbatch protocols do not readily allow for screening at anything typically other than atmospheric pressure. The rigidity of the PDMS matrix enables experiments to probe the effects of pressure on target crystallization on-chip.

The surface on which the crystallization ‘drop’ sits can affect experimental success and crystal quality. Examples of solid support contact surfaces used in vapor diffusion and microbatch protocols include either polystyrene or silanized glass. Both types of supports can show different propensities to promote or inhibit crystal growth, depending on the target. In addition, the crystallization ‘drop’ is in contact with either air or some type of poly-carbon oil, depending on whether the experiment is a vapor-diffusion or microbatch setup, respectively. Air contact has the disadvantage in that free oxygen reacts readily with biological targets, which can lead to protein denaturation and inhibit or degrade crystallization success. Oil allows trace hydrocarbons to leach into the crystallization experiment, and can similarly inhibit or degrade crystallization success.

Microfluidics device designs in accordance with embodiments of the present invention may overcome these limitations by providing a nonreactive, biocompatible environment that completely surrounds the crystallization reaction. Moreover, the composition of the crystallization chambers in the microfluidics chips can conceivably be varied to provide new surfaces for contacting the crystallization reaction; this would allow for routine screening of different surfaces and surface properties to promote crystallization.

Crystallization targets, particularly those of biological origin, may often be modified to enable crystallization. Such modifications include but are not limited to truncations, limited proteolytic digests, site-directed mutants, inhibited or activated states, chemical modification or derivatization, etc. Target modifications can be time consuming and costly; modified targets require the same thorough screening as do unmodified targets. Microfluidics devices of the present invention work with such modified targets as readily as with the original target, and provide the same benefits.

The effect of gravity as a parameter for crystallization is yet another understudied crystallization parameter, because of the difficulty of varying such a physical property. Nonetheless, crystallization experiments of biological samples in zero gravity environments have resulted in the growth of crystals of superior quality than those obtained on Earth under the influence of gravity.

The absence of gravity presents problems for traditional vapor-diffusion and microbatch setups, because all fluids must be held in place by surface tension. The need to often set up such experiments by hand also poses difficulties because of the expense of maintaining personnel in space. Microfluidics devices in accordance with embodiments of the present invention, however, would enable further exploration of microgravity as a crystallization condition. A compact, automated metering and crystal growth system would allow for: 1) launching of satellite factory containing target molecules in a cooled, but liquid state, 2) distribution of targets and growth of crystals, 3) harvesting and cryofreezing of resultant crystals, and 4) return of cryo-stored crystals to land-based stations for analysis.

Embodiments of crystallization structures and methods in accordance with the present invention offer a number of advantages over conventional approaches. One advantage is that the extremely small volumes (nanoliter/sub-nanoliter) of sample and crystallizing agent permit a wide variety of recrystallization conditions to be employed utilizing a relatively small amount of sample.

Another advantage of crystallization structures and methods in accordance with embodiments of the present invention is that the small size of the crystallization chambers allows crystallization attempts under hundreds or even thousands of different sets of conditions to be performed simultaneously. The small volumes of sample and crystallizing agent employed in crystallization also result in a minimum waste of valuable purified target material.

A further advantage of crystallization in accordance with embodiments of the present invention is relative simplicity of operation. Specifically, control over flow utilizing parallel actuation requires the presence of only a few control lines, with the introduction of sample and crystallizing agent automatically performed by operation of the microfabricated device permits very rapid preparation times for a large number of samples.

Still another advantage of crystallization systems in accordance with embodiments of the present invention is the ability to control solution equilibration rates. Crystal growth is often very slow, and no crystals will be formed if the solution rapidly passes through an optimal concentration on the way to equilibrium. It may therefore be advantageous to control the rate of equilibration and thereby promote crystal growth at intermediate concentrations. In conventional approaches to crystallization, slow-paced equilibrium is achieved using such techniques as vapor diffusion, slow dialysis, and very small physical interfaces.

Another potential advantage of the creation and use of microfluidic free interfaces in accordance with embodiments of the present invention is the ability to preserve intact crystals that have formed as a result of the diffusive mixing. Specifically, once a crystal has nucleated and grown, it is frequently desirable to reduce the temperature of the crystal in order to preserve its structure during subsequent x-ray analysis. Such analysis may induce radiation damage, which undesirably decays or destroys the crystal.

Accordingly, it is desirable to expose the crystal to a cryogen or other low temperature compound to preserve the crystal intact. The exclusively diffusive mixing allowed by embodiments of microfluidic free interfaces in accordance with embodiments of the present invention allow for the exposure of such cryogens to the crystal, such that the reduction in temperature does not result in damage or change in the crystal structure. For example, in certain embodiments in accordance with the present invention, once the presence of a crystal has been identified, a second microfluidic free interface between the crystal-containing solution and a cryogenic solution can be created. The cryogenic contents of that chamber can then be allowed to diffuse into the crystallization chamber. In this scheme, large number of cryogenic types and concentrations can readily be screened without changing the location of the crystal in-situ.

2. Solubility Studies (Proteomics)

The above discussion has focused upon protein crystallization applications wherein the environment of the solvent containing a protein sample was altered by diffusion of a crystallizing agent across a microfluidic free interface, resulting in a change in phase of the protein and the formation of a crystal. However, an important precursor to such a crystallization study is identifying and obtaining a solution in which the protein is sufficiently soluble and mono-disperse to form a crystal upon exposure to a crystallizing agent.

Accordingly, an alternative application for microfluidic free interface diffusion in accordance with an embodiment of the present invention is to determine solubility of a protein in various solvents. Specifically, a free microfluidic interface could be utilized to meter the flow of a solvent to a protein sample, and thus to identify the solubility characteristic of the protein. Specifically, the presence of a free microfluidic interface would create a uniform and precise concentration gradient along a channel of known volume, thereby enabling a researcher to identify with precision the ability of a known amount of protein sample entering into solution. This information could in turn be utilized to allow for the creation of concentrated protein solutions that are in turn susceptible to forming crystals.

3. Sample Purification Through Quasi-Dialysis

The above discussion has focused upon protein crystallization and proteomics solubility applications wherein the relative amount of sample and solvent are precisely controlled to govern concentration of the sample in solution, and hence the propensity of the sample to emerge from solution in solid, crystalline form. However, a number of factors other than sample concentration may also play an important role in sample behavior.

One such factor is sample purity. One artifact of many common processes for protein isolation and/or purification is the concentration of small salts with the sample. These salts can stabilize a protein in solution, and otherwise interfere with crystal nucleation and growth. One conventional approach for removing these salts is through dialysis across a membrane. However, this approach typically takes place on the macroscopic scale, involving the use of relatively large sample volumes and involving the coarse removal of bulk quantities of salts.

Thus yet another potential alternative application for microfluidic free interface diffusion in accordance with an embodiment of the present invention is in the removal of unwanted components from a sample solution. For example, a solution comprising isolated and/or purified protein that contains a significant concentration of small salts may be positioned in a microfluidic channel on one side of a valve. Upon deactuation of the valve, a microfluidic free interface between the protein solution and a diluent may be created, such that the smaller salts rapidly diffuse into the diluent and are thereby depleted in the sample. Of course, some amount of protein present in the sample would also be expected to diffuse across the microfluidic free interface and thereby be lost to the diluent during this process. However, disparity in size between the small salt and the large protein molecule would be expected to constrain relative diffusion rates to limit the loss of protein in the sample.

4. Growth of Other than Protein Crystals

Embodiments in accordance with the present invention have focused thus far on the growth of crystals biological macromolecules such as proteins. However, other types of molecules also form crystals whose three dimensional structure may be studied. One such promising candidate for crystal formation are inorganic nanocrystals which may be utilized to entrap single atoms.

Recently the scientific community has exhibited interest in growing nano-crystals from inorganic materials. Such crystals have dimensions on the nanometer scale, and in some instances are able to enclose individual atoms. These crystals exhibit optical properties such as absorption and emission of light, which are governed by quantum mechanical effects. These optical properties may be controlled or varied according to such factors as the size and quality of the crystal, and the identity of the atomic species enclosed thereby.

As with a protein crystal, the size and quality of such a nano-crystal depends to a large degree upon conditions in which it is grown. Therefore, it may prove important to control both the phase-space evolution, and the rate of this evolution during growth of a nano-crystal. Such control can be accomplished utilizing microfluidic free interface diffusion in accordance with embodiments of the present invention, as has been described previously in connection with protein crystallization.

Moreover, in certain cases it may be prove of interest to determine the effect upon crystal properties of changing the concentration of one element of the crystallization. This may be accomplished using microfluidic free interface diffusion in accordance with embodiments of the present invention by setting up a gradient through a string of connected chambers, with a large reservoir at either end (or simply along a channel connecting two large reservoirs). Crystals are then grown at each condition (or along a continuum of conditions) and then are interrogated optically to determine their absorption/emission properties. In this manner, it is possible to independently vary a single crystallization parameter (such as the concentration of a sample, a crystallizing agent, a precipitant, or some other species).

It is also possible to grow on chip a spatial distribution of nanocrystals with different optical properties.

5. Diffusive Immunoassays

Another potentially valuable application for microfluidic free interfaces in accordance with embodiments of the present invention is analysis of binding of analytes to targets. Conventionally, binding of an analyte to a target has been determined through techniques such as chromatography, wherein the changed velocity of an analyte/target combination relative to either of the analyte or target alone in a convective flow is employed to identify the presence of the bound combination.

Embodiments in accordance with the present invention exploit the fact that the binding of analyte to a target will also result in a changed size and hence coefficient of diffusion of the analyte/target combination relative to either of the analyte or target alone. This changed diffusion coefficient can be utilized to identify the bound combination.

FIG. 45A shows a simplified plan view of one embodiment of a device that is able to detect diffusion across a microfluidic free interface in accordance with the present invention. Chambers 8000 and 8002 containing fluids A and B respectively are connected by channel 8004 having a width that is restricted in at least one dimension. Valve 8006, which is initially closed, is positioned within channel 8004 to maintain fluids A and B separate. Chamber 8003 is in fluid communication with chamber 8002 through a connecting channel 8005 and also contains fluid B.

First fluid A contains an analyte 8008 detectable through analytical techniques, including but not limited to detecting changes in fluoresce, refractive index, conductivity, light-scattering, and the use of colorimetric sensors. Second fluid B contains a target 8010 that is known to bind to analyte 8008. Upon binding of analyte 8008 to target 8010, the analyte/target combination exhibit a substantially reduced coefficient of diffusion relative to that of the analyte alone.

At an initial time T₀, valve 8006 is opened to create a microfluidic free interface 8012 between fluids A and B. Commencing at time T₁, analyte 8008 begins to diffuse through channel 8004, chamber 8002, and channel 8005 into chamber 8003. Absent a binding reaction between the analyte 8008 and target 8010, the analyte would be expected to begin appearing in end chamber 8003 at a given rate over time. This is shown in FIG. 45B, which plots intensity of a signal expected to be detected in the end chamber 8003 versus time of diffusion.

However, because some reaction occurs in the channel between analyte 8008 and target 8010, some percentage of analyte 8008 binds to target 8010, effectively slowing the rate of diffusion of this percentage of the analyte. As a result, the observed intensity of the signal in the end chamber over time will be reduced, reflecting the slower rate of diffusion of the bound analyte/target combination. This is also shown in FIG. 45B, which plots intensity of an observed signal in the second chamber versus time of diffusion.

The difference between expected and observed signal intensity over time may be analyzed to reveal potentially valuable information regarding the fluidic system. For example, where the concentration of target and analyte in the respective solutions is known, analysis of evolution over time of the intensity signal can indicate the reactivity between the target and analyte, i.e. the rate at which the analyte binds to the target and whose diffusion is slowed as result. Alternatively, where reactivity between the target and the analyte is known, analysis of the time evolution of the intensity signal can indicate the initial concentration of the target. Further alternatively, where initial concentrations and reactivity are known, analysis of time evolution of the intensity signal can indicate the size of the target, as manifested by a reduced diffusion rate of the bound analyte/target combination versus the analyte alone.

The example just provided utilized an analysis of the time evolution of intensity of a signal from a diffusing analyte. However, embodiments in accordance with the present invention are not limited to this particular approach.

For example, FIG. 46 shows an alternative embodiment of a structure for use in diffusive assays. Structure 8100 comprises microfluidic channel 8102 of constant width and depth that is bisected by valve 8104. Fluid A containing detectable analyte 8106 is positioned on one side of valve 8104. Fluid B containing target 8108 is positioned on the other side of valve 8104.

At an initial time T₀, valve 8104 opened to create a microfluidic free interface 8110 between fluids A and B. Commencing at time T₁, analyte 8106 begins to diffuse across interface 8110. Absent a binding reaction between the analyte 8106 and target-8108, after a certain time the profile of intensity of the signal over the distance of the channel would be expected to exhibit a characteristic shape. This is shown in FIG. 46B, which plots intensity of a signal expected to be detected in the second chamber, versus distance of diffusion.

However, because some reaction occurs between analyte 8106 and target 8108, some percentage of the analyte binds to the target, effectively slowing the rate of diffusion of this percentage of the analyte. As a result, the observed intensity of the signal in the second half of the channel would be changed, reflecting the slower rate of diffusion of the bound analyte/target combination. This is also shown in FIG. 46B, which plots intensity of an observed signal in the second chamber versus distance of diffusion.

As with the embodiment shown in FIGS. 45A-B, the difference between expected and observed signal intensity over distance may be analyzed to reveal potentially valuable information. For example, where the concentration of target and analyte in the respective solutions is known, analysis of spatial evolution of the intensity signal can indicate the reactivity between the target and analyte, i.e. the rate at which the analyte binds to the target and whose diffusion is slowed as result. Alternatively, where reactivity between the target and the analyte is known, analysis of spatial evolution of the intensity signal can indicate the initial concentration of the target. Further alternatively, where initial concentrations and reactivity are know, analysis of spatial evolution of the intensity signal can indicate the mass of the target, as manifested by reduction in the rate of diffusion of the bound analyte/target combination versus the analyte alone.

6. Viscosity Measurement

The above discussion focused upon identifying properties of a chemical sample such as concentration or reactivity based upon diffusion of the sample across a microfluidic free interface under known conditions. In accordance with alternative embodiments of the present invention however, properties of the fluid in which diffusion is taking place, rather than of the sample diffusing within the fluid, may be determined by observing the character of diffusion across a free microfluidic interface.

For example, one conventional approach to measuring viscosity calls for inserting a piston into a closed vessel containing the sample fluid, and then measuring the torque required to rotate the piston in the vessel. While this approach is adequate for measuring the viscosity of larger fluid samples, it offers the disadvantage of requiring and consuming relatively large volumes of sample fluid. Such volumes may not be available for analysis, particularly in the context of biological analysis.

In accordance with an alternative embodiment of the present invention, diffusion of a marker of known size and diffusion coefficient (i.e. a fluorescently labeled bead or macromolecule) across a microfluidic free interface created within the fluid may be utilized to determine its viscosity. Such a technique is particularly applicable to analysis of the viscosity of biological or physiological samples, as the dimensions of the microfluidic channels in which diffusion occurs occupies relatively small volumes.

7. Competitive Binding Assay

A different type of assay from the diffusive immunoassay discussed above is the competitive binding assay. A competitive binding assay determines the propensity of a competitor ligand to displace an existing ligand bound to a target molecule such as a protein. This displacement tendency may be determined by mixing the original ligand/target molecule combination with the competitor ligand at various concentrations. Conventionally, such binding assays require a large number of separate experiments to cover a suitable range of concentrations at sufficient resolution.

However, diffusion across a microfluidic free interface in accordance with an embodiment of the present invention allows for establishing continuous or discrete gradients such that a large number of assay conditions can be sampled at once. For example, in accordance with one embodiment of the present invention, a competitive binding assay may be performed utilizing a discrete gradient of competitor ligand concentrations created in a string of chambers connected by channels.

This approach is illustrated in FIG. 49, which shows a string of chambers 8300 a-c connected by narrow channels 8302 and positioned between reservoirs 8304 and 8306. Reservoir 8304 contains a high concentration the competitor ligand 8308, and reservoir 8306 does not contain the competitor ligand. Flow into each of chambers 8300 a-c from the reservoirs or from each other may be independently controlled by valves 8305 located in connecting channels 8302. Concentrations at either end of the string of chambers may be fixed by the large volumes of the reservoirs relative to the chambers, or by continuous flows of material to the reservoirs.

Initially, the chambers in the string are charged with a first fluid containing a certain concentration of the original ligand 8310 bound to target 8312. Next, chambers 8300 a-c are placed in fluid communication with the reservoirs and with each other by opening interface valves 8305. Once these interface valves are opened, a microfluidic free interface is established between the chambers, and the solutions equilibrate by diffusion. The position of each chamber along the string determines the concentration of the competitor ligand in that particular chamber.

The competitor ligand is typically much smaller than the target molecule and therefore equilibrates quickly. The original ligand displaced from the target by the competitive ligand is free to diffuse from the target chamber to another chamber. Owing to reduced size of the original ligand relative to the ligand/target combination, the displaced original ligand would be expected to exhibit a greater diffusion coefficient and hence diffuse at a faster rate. In a manner similar to that described above in connection with FIGS. 45A-B and 46A-B, if the original ligand is fluorescent, its presence and hence the rate of displacement from the target may be revealed by analyzing fluorescence signals for deviation from expected temporal or spatial profiles.

While the embodiment just described monitors changes in location of a fluorescent signal over time to identify competitive binding, in accordance with alternative embodiments of the present invention, a diminution in intensity of a fluorescent signal may also be detected to perform competitive binding assays. Specifically, fluorescent ligands may be utilized in the presence of quenching molecules. A fluorescent molecule is quenched when it interacts another molecule to suppress fluorescent emission.

FIG. 50 shows a plan view of a simplified embodiment of a microfluidic system for performing competitive binding assays utilizing a quenching material. First chamber 8400 contains target molecule 8402, first fluorescent ligand 8404, and quenching molecule 8406. First chamber 8400 is connected via a microfabricated channel 8408 to a second chamber 8410 containing competitor ligand 8412, and the same concentrations of target molecule 8402, first fluorescent ligand 8404, and quenching molecule 8406 as are present in first chamber 8400.

When first fluorescent ligand 8404 is bound to target 8402, it cannot interact with the quenching molecule and thus a fluorescent emission may be observed. When the first fluorescent ligand is not bound to the target, however, it is quenched and therefore no fluorescent emission is observed. In this manner, a decline in fluorescent signal intensity may indicate competitive binding behavior.

Valve 8414 originally separating the contents of chambers 8400 and 8410 is opened, creating microfluidic free interface 8414 and allowing the two fluids to mix by diffusion. Since the fluids are identical except for the presence of the competitor ligand in the second chamber, only the competitor ligand experiences net diffusive transport between the chambers and establishes a linear gradient between the chambers. By observing the fluorescent intensity along different portions of the channel, the concentration of competitor ligand necessary to displace the first fluorescent ligand can be determined. Alternatively, the level of fluorescence in the end chambers over time can be monitored to determine concentration of the competitor ligand necessary to displace the first fluorescent ligand from the target. Using either method, a single experiment may be employed to determine the outcome of the competitive binding assay under a continuum of conditions.

8. Substrate/Inhibitor Assays

Activity of an enzyme can be measured in response to the presence of different concentrations of a ligand. For example, a protein catalyzing a reaction may be impaired by a small molecule that binds strongly to the active site of the protein. A substrate turnover assay may be used to determine the effect of such a ligand on enzyme activity.

Where the enzyme catalyses a reaction that creates a fluorescent product, the activity of the enzyme may be determined by monitoring fluorescence. One example of catalysis of a fluorescence-producing reaction is the hydrolyzation of Beta-D-galactopyranoside into resorufin by the enzyme Beta-galactosidase. Resorufin is a fluorescent product that may be monitored over time. Ramsey et. al, Analytical Chem. 69, 3407-3412 (1997). By determining activity of the enzyme as a function of the concentration of an inhibiting molecule, the Michaelis-Menten rate constants of the reaction can be determined. The generation of a diffusion gradient across a microfluidic free interface in accordance with an embodiment of the present invention can enable determination of this concentration dependence in a single experiment.

FIG. 51 shows a simplified plan view of an embodiment of a microfluidic structure for performing such a substrate turnover assay. Again, reservoirs 8500 and 8502 are connected by channel 8504 having a width that is much smaller than the dimensions of the chambers. Reservoirs 8500 and 8502 contain substrate 8509 at the same concentration, and reservoir 8502 also contains inhibitor molecule 8508. A gradient of inhibitor molecule concentration is achieved by maintaining a fixed concentration of material in the reservoirs 8500 and 8502, due either to their large volumes relative to the channel, or to replenishment of material from an external source. Chambers may or may not be included along the length of channel 8504 to provide larger volumes for reaction sites at particular concentrations.

Once the concentration gradient of inhibitor molecule has been established, valves 8510 are actuated to isolate sections of channel (or individual chambers connected by the channel) that are subsequently mixed with separate wells 8512 containing enzyme 8514. By monitoring fluorescence of the each well over time, it is possible to determine the activity the enzyme in each chamber.

Detection of activity in accordance with embodiments of the present invention is not required to be through fluorescence. Depending upon the substrate to be identified, enzyme activity and the resulting catalysis could be determined through analysis of colorimetry, index of refraction, surface plasmon resonance, or conductivity, to name only a few.

9. Bioreactor

Only a small percentage of known microorganisms can be cultured and grown in a laboratory setting. This relatively low success rate is largely attributable to the substantial number of factors typically required to induce cell growth. The success of a structure for promoting cell growth depends fundamentally on the ability to continuously supply nutrients to the growing cells and to remove waste products that are created by the cell metabolism.

In accordance with embodiments of the present invention, microfluidic free interface diffusion could be employed to continuously supply cells with needed food, and to eliminate waste products. Cells could be introduced and then contained in chambers that are connected to various reservoirs containing food, growth factors, buffer, and other materials to support cell growth, via channels that are too small to let the cells pass. Free interface diffusion could then be used to control the flux of food, growth factors, or waste to and from the chambers.

For example, FIGS. 27A and 27B show plan and cross-sectional views (along line 45B-45B′) respectively, of one embodiment of a cell cage structure in accordance with the present invention. Cell cage 4500 is formed as an enlarged portion 4500 a of a flow channel 4501 in an elastomeric block 4503 in contact with substrate 4505. Cell cage 4500 is similar to an individual cell pen as described above in FIGS. 26A-D, except that ends 4500 b and 4500 c of cell cage 4500 do not completely enclose interior region 4500 a. Rather, ends 4500 a and 4500 b of cage 4500 are formed by a plurality of retractable pillars 4502. Pillars 4502 may be part of a membrane structure of a normally-closed valve structure as described extensively above in connection with FIGS. 21A-J.

Specifically, control channel 4504 overlies pillars 4502. When the pressure in control channel 4504 is reduced, elastomeric pillars 4502 are drawn upward into control channel 4504, thereby opening end 4500 b of cell cage 4500 and permitting a cell to enter. Upon elevation of pressure in control channel 4504, pillars 4502 relax downward against substrate 4505 and prevent a cell from exiting cage 4500.

Elastomeric pillars 4502 are of a sufficient size and number to prevent movement of a cell out of cage 4500, but also include gaps 4508 which allow the flow or diffusion of nutrients into cage interior 4500 a in order to sustain cell(s) stored therein. Pillars 4502 on opposite end 4500 c are similarly configured beneath second control channel 4506 to permit opening of the cage and removal of the cell as desired.

Microfluidic free interface diffusion in accordance with embodiments of the present invention could be utilized to provide a steady supply of nutrients and growth factors to cells growing within the cell cage, and to remove waste products such as low molecular weight salts from the cell cage. One embodiment is shown in FIG. 48, wherein cell cage 8200 is in fluid communication with dead-ended branch channel 8202 of flow channel 8204, which supplies nutrients 8205 by diffusion across the microfluidic free interface 8206. Waste products 8207 from cell cage 8200 in turn also diffuse across microfluidic free interface 8206 and enter into flow channel 8204 for removal.

Control over the rate of diffusion of materials to and from the cell cage could be accomplished by varying microfluidic channel dimensions such as width, height, and length, the number of connecting channels, the size of reservoirs, the size of chambers, and the concentration of nutrient supplies and or waste products. By varying the above-listed factors, amongst others, it is possible to regulate the rate of cell growth and division, and hence the maximum density of cells. And because at microfluidic dimensions there may be no appreciable concentration gradient within larger volume chambers connected by a narrow channel, the entire cell population within such a chamber may be exposed simultaneously to similar conditions.

Utilizing microfluidic free interface diffusion in accordance with embodiments of the present invention, it may also be possible to culture two or more different strains of cells positioned in adjacent chambers, while allowing for diffusion of nutrients, waste, and other products of cell growth to move between the chambers. This type of microfluidic structure and method may be particularly valuable in studying interactions between proximate cells, including but not limited to cell signaling, cell/cell toxicity, and cell/cell symbiosis.

Finally, microfluidic free interface diffusion in accordance with embodiments of the present invention could further be used to establish gradients of nutrient/waste conditions in order to screen for desired cell growth conditions. The use of diffusion for growing cell cultures is practical at these small dimensions, where diffusion lengths are short and volumes are sufficiently small to allow diffusive transport to significantly change conditions over a short period of time.

10. Chemotaxis

Still another potential application for embodiments in accordance with the present invention is the study of chemotaxis. Chemotaxis is the change in direction of movement of a motile cell in response to a concentration gradient of a specific chemical. This change in the direction may result in the cell moving toward or away from the sensed chemical at a specific rate. Chemotaxis plays a key role in immuno-responses, wherein components of the immune system migrate to and destroy foreign agents in response to a sensed change in chemical environment attributable to the presence of the foreign agents.

Microfluidic free interfaces created accordance with embodiments of the present invention may provide a valuable way to precisely reproduce chemical concentration gradients. The activity of motile cells in response to exposure to these concentration gradients would enable researchers to study and understand the phenomenon of chemotaxis.

11. Creating a Polymer with Cross-Linking/Porosity Gradient

In many biological applications, polymerized gels are used to separate molecules based on their differing mobility through the gel, which is typically of uniform composition. However, the ability of the get to efficiently separate molecules may be improved by imposing a gradient on properties such as pH, density, or polymerization cross-linking during formation of the gel.

Thus in accordance with an embodiment of the present invention, microfluidic free interface diffusion may be used to fabricate polymerized gels having well defined gradients of particular properties. For instance, in one example a concentrated, unpolymerized, photo-curable gel (such as agarose) may be introduced into a microfluidic chamber in fluid communication via a microfluidic channel with a second microfluidic chamber containing a buffer solution. As the contents of the respective chambers diffuse into one another through the channel, a gel concentration gradient is established. Once this gradient has formed, the gel may be exposed to UV radiation to cause cross-linking creating a solid gel. The gel concentration gradient, and hence the porosity of the gel, is fixed into the gel as a result of this polymerization.

A gel exhibiting a gradient as just described may be used to determine the size of a macromolecule by applying a potential and allowing the molecule to migrate through the gel from the area of greatest porosity and lowest gel concentration, to the area of least porosity and highest gel concentration. Once the average pore size of the gel becomes too small to allow the molecule to pass, the molecule will remain trapped in the gel and detectable by standard fluorescent or silver staining methods. The location of the molecules in the gel would therefore pose a direct reflection of the size of the molecules.

In an alternative embodiment of forming gels with superimposed gradients using microfluidic free interface diffusion in accordance with the present invention, the polymerized gel may be formed with an imprinted pH gradient. In such an alternative embodiment, a first chamber is filled with an unpolymerized gel with one buffer at one pH and a second chamber is filled with a combination of the unpolymerized gel at the same concentration with a second buffer at a second pH. The molecules are then allowed to diffuse across the microfluidic channel connecting the chambers to establish a pH gradient. The gel is then polymerized creating a solid gel of uniform porosity. Where the diffusing molecule is capable of being permanently bonded to the gel matrix, a uniform pH gradient may be permanently established along the length of the connecting channel. Such a pH gradient may be employed in combination with an applied electric field to separate molecules moving through the gel based upon their different isoelectric points.

12. Drug Delivery Systems

The ability to control the rate of flux out of or into a microfluidic chamber enables the timed release of chemicals or drugs, without the need for active control strategies such as clocks. For example, an implanted microfluidic drug delivery device could contain a plurality of chambers connected to an exit orifice through microfluidic channels of varying geometries. In this way the rate of transport of each of the chemicals in the chambers could be controlled separately and in combination. Utilizing such a structure, several different drugs could be introduced to the body at different and well defined rates.

For example, in an embodiment of a drug delivery device utilizing microfluidic free interface diffusion, a plurality of chambers may first be connected via channels of varying dimensions to chambers filled with a buffer solution. These buffer-filled chambers may then in turn be connected via channels of various dimensions to an exit orifice. Dimensions of the microfluidic chambers and connecting channels would allow control over both the rate and timing of release of the drugs into the body from the chambers. By judicious selection of channel dimensions such as width and length, and chamber volumes and shapes, and drug concentrations, it is possible to achieve very precise control over the total rate of drug delivery over time.

13. Mixing of Highly Reactive Chemical Species

Still another example of a potential application for microfluidic free interfaces in accordance with embodiments of the present invention is in the controlled mixing of highly reactive species. For example, under certain circumstances it may be desirable to cause a reaction to occur at a slow or gradual pace between two or more chemical species. Under conventional conditions involving convective flow however, the species may react suddenly and/or violently, mitigating the usefulness of the reaction products.

In accordance with embodiments of the present invention, however, contact and reaction between the chemical species is governed exclusively by diffusion with no convective flow. Therefore it may be possible to conduct a reaction between the chemical species at a slowed and governable pace. One field wherein embodiments in accordance with the present invention could offer a considerable advantage is in the preparation of explosives or other potentially unstable products, wherein rapidly changing reactant concentrations could give rise to hazardous or otherwise undesirable conditions. Embodiments in accordance with the present invention could also be potentially useful for combinitoric synthesis applications where an end or intermediate product is unstable, highly reactive, and/or volatile.

14. Biological Assays

Embodiments of microfluidic fee interfaces in accordance with the present invention may be employed for a variety of applications. Examples of such applications are summarized below. A more complete description of possible applications may be found in PCT application PCT/US01/44869, filed Nov. 16, 2001 and entitled “Cell Assays and High Throughput Screening”, hereby incorporated by reference for all purposes. Examples of microfluidic structures suitable for performing such applications include those described herein, as well as others described in U.S. patent application Ser. No. 10/118,466, “Nucleic Acid Amplification Utilizing Microfluidic Devices”, filed Apr. 5, 2002, hereby incorporated by reference for all purposes.

A wide variety of binding assays can be conducted utilizing the microfluidic devices disclosed herein. Interactions between essentially any ligand and antiligand can be detected. Examples of ligand/antiligand binding interactions that can be investigated include, but are not limited to, enzyme/ligand interactions (e.g., substrates, cofactors, inhibitors); receptor/ligand; antigen/antibody; protein/protein (homophilic/heterophilic interactions); protein/nucleic acid; DNA/DNA; and DNA/RNA. Thus, the assays can be used to identify agonists and antagonists to receptors of interest, to identify ligands able to bind receptors and trigger an intracellular signal cascade, and to identify complementary nucleic acids, for example. Assays can be conducted in direct binding formats in which a ligand and putative antiligand are contacted with one another or in competitive binding formats well known to those of ordinary skill in the art.

Heterogenous binding assays involve a step in which complexes are separated from unreacted agents so that labeled complexes can be distinguished from uncomplexed labeled reactants. Often this is achieved by attaching either the ligand or antiligand to a support. After ligands and antiligands have been brought into contact, uncomplexed reactants are washed away and the remaining complexes subsequently detected.

The binding assays conducted with the microfluidic devices provided herein can also be conducted in homogeneous formats. In the homogeneous formats, ligands and antiligands are contacted with one another in solution and binding complexes detected without having to remove uncomplexed ligands and antiligands. Two approaches frequently utilized to conduct homogenous assays are fluorescence polarization (FP) and FRET assays.

The microfluidic devices can also be utilized in a competitive formats to identify agents that inhibit the interaction between known binding partners. Such methods generally involve preparing a reaction mixture containing the binding partners under conditions and for a time sufficient to allow the binding partners to interact and form a complex. In order to test a compound for inhibitory activity, the reaction mixture is prepared in the presence (test reaction mixture) and absence (control reaction mixture) of the test compound. Formation of complexes between binding partners is then detected, typically by detecting a label borne by one or both of the binding partners. The formation of more complexes in the control reaction then in the test reaction mixture at a level that constitutes a statistically significant difference indicates that the test compound interferes with the interaction between the binding partners.

Immunological assays are one general category of assays that can be performed with the microfluidic devices in accordance with embodiments of the present invention. Certain assays are conducted to screen a population of antibodies for those that can specifically bind to a particular antigen of interest. In such assays, a test antibody or population of antibodies is contacted with the antigen. Typically, the antigen is attached to a solid support. Examples of immunological assays include enzyme linked immunosorbent assays (ELISA) and competitive assays as are known in the art.

Utilizing the microfluidic devices provided herein, a variety of enzymatic assays can be performed. Such enzymatic assays generally involve introducing an assay mixture containing the necessary components to conduct an assay into the various branch flow channels. The assay mixtures typically contain the substrate(s) for the enzyme, necessary cofactors (e.g., metal ions, NADH, NAPDH), and buffer, for example. If a coupled assay is to be performed, the assay solution will also generally contain the enzyme, substrate(s) and cofactors necessary for the enzymatic couple.

Microfluidic devices in accordance with embodiments of the present invention can be arranged to include a material that selectively binds to an enzymatic product that is produced. In some instances, the material has specific binding affinity for the reaction product itself. Somewhat more complicated systems can be developed for enzymes that catalyze transfer reactions. Certain assays of this type, for example, involve incubating an enzyme that catalyzes the transfer of a detectable moiety from a donor substrate to an acceptor substrate that bears an affinity label to produce a product bearing both the detectable moiety and the affinity label. This product can be captured by material that includes a complementary agent that specifically binds to the affinity label. This material typically is located in a detection region such that captured product can be readily detected. In certain assays, the material is coated to the interior channel walls of the detection section; alternatively, the material can be a support located in the detection region that is coated with the agent.

Certain assays utilizing the present devices are conducted with vesicles rather than cells. Once example of such an assay is a G-protein coupled receptor assay utilizing fluorescent correlation spectroscopy (FCS). Membrane vesicles constructed from cells that over-express the receptor of interest are introduced into a main flow channel. Vesicles can either be premixed with inhibitor and introduced via branch flow channels or via one of the main flow channels prior to being mixed with a fluorescent natural ligand which is also introduced by a main flow channel. Components are allowed to incubate for the desired time and fluorescent signals may be analyzed directly in the flow chamber using an FCS reader such as the Evotec/Zeiss Confocor (a single or dual photon counting device).

FRET assays can also be utilized to conduct a number of ligand-receptor interactions using the devices disclosed herein. For example, a FRET peptide reporter can be constructed by introducing a linker sequence (corresponding to an inducible domain of a protein such as a phosphorylation site) into a vector encoding for a fluorescent protein composed of blue- and red-shifted GFP variants. The vector can be a bacterial (for biochemical studies) or a mammalian expression vector (for in vivo studies).

Assays of nuclear receptors can also be performed with the present microfluidic devices. For example, FRET-based assays for co-activator/nuclear receptor interaction can be performed. As a specific example, such assays can be conducted to detect FRET interactions between: (a) a ligand binding domain of a receptor tagged with CFP (cyan fluorescent protein, a GFP derivative), and (b) a receptor binding protein (a coactivator) tagged with the Yellow fluorescent protein (YFP).

Fluorescence polarization (FP) can be utilized to develop high throughput screening (HTS) assays for nuclear receptor-ligand displacement and kinase inhibition. Because FP is a solution-based, homogeneous technique, there is no requirement for immobilization or separation of reaction components. In general, the methods involve using competition between a fluorescently labeled ligand for the receptor and related test compounds.

A number of different cell reporter assays can be conducted with the provided microfluidic devices. One common type of reporter assay that can be conducted include those designed to identify agents that can bind to a cellular receptor and trigger the activation of an intracellular signal or signal cascade that activates transcription of a reporter construct. Such assays are useful for identifying compounds that can activate expression of a gene of interest. Two-hybrid assays, discussed below, are another major group of cell reporter assays that can be performed with the devices. The two-hybrid assays are useful for investigating binding interactions between proteins.

Often cell reporter assays are utilized to screen libraries of compounds. In general such methods involve introducing the cells into the main flow channel so that cells are retained in the chambers located at the intersection between the main flow channel and branch channels. Different test agents (e.g., from a library) can then be introduced into the different branch channels where they become mixed with the cells in the chambers. Alternatively, cells can be introduced via the main flow channel and then transferred into the branch channel, where the cells are stored in the holding areas. Meanwhile, different test compounds are introduced into the different branch flow channels, usually to at least partially fill the chambers located at the intersection of the main and branch flow channels. The cells retained in the holding area can be released by opening the appropriate valves and the cells transferred to the chambers for interaction with the different test compounds. Once the cells and test compounds have been mixed, the resulting solution is returned to the holding space or transported to the detection section for detection of reporter expression. The cells and test agents can optionally be further mixed and incubated using mixers of the design set forth above.

Cells utilized in screening compounds to identify those able to trigger gene expression typically express a receptor of interest and harbor a heterologous reporter construct. The receptor is one which activates transcription of a gene upon binding of a ligand to the receptor. The reporter construct is usually a vector that includes a transcriptional control element and a reporter gene operably linked thereto. The transcriptional control element is a genetic element that is responsive to an intracellular signal (e.g., a transcription factor) generated upon binding of a ligand to the receptor under investigation. The reporter gene encodes a detectable transcriptional or translational product. Often the reporter (e.g., an enzyme) can generate an optical signal that can be detected by a detector associated with a microfluidic device.

A wide variety of receptor types can be screened. The receptors often are cell-surface receptors, but intracellular receptors can also be investigated provided the test compounds being screened are able to enter into the cell. Examples of receptors that can be investigated include, but are not limited to, ion channels (e.g., calcium, sodium, potassium channels), voltage-gated ion channels, ligand-gated ion channels (e.g., acetyl choline receptors, and GABA (gamma-aminobutyric acid) receptors), growth factor receptors, muscarinic receptors, glutamate receptors, adrenergic receptors, dopamine receptors.

Another general category of cell assays that can be performed is the two hybrid assays. In general, the two-hybrid assays exploit the fact that many eukaryotic transcription factors include a distinct DNA-binding domain and a distinct transcriptional activation domain to detect interactions between two different hybrid or fusion proteins. Thus, the cells utilized in two-hybrid assays include the construct(s) that encode for the two fusion proteins. These two domains are fused to separate binding proteins potentially capable of interacting with one another under certain conditions. The cells utilized in conducting two-hybrid assays contain a reporter gene whose expression depends upon either an interaction, or lack of interaction, between the two fusion proteins.

In addition to the assays just described, a variety of methods to assay for cell membrane potential can be conducted with the microfluidic devices disclosed herein. In general, methods for monitoring membrane potential and ion channel activity can be measured using two alternate methods. One general approach is to use fluorescent ion shelters to measure bulk changes in ion concentrations inside cells. The second general approach is to use of FRET dyes sensitive to membrane potential.

The microfluidic devices disclosed herein can be utilized to conduct a variety of different assays to monitor cell proliferation. Such assays can be utilized in a variety of different studies. For example, the cell proliferation assays can be utilized in toxicological analyses, for example. Cell proliferation assays also have value in screening compounds for the treatment of various cell proliferation disorders including tumors.

The microfluidic devices disclosed herein can be utilized to perform a variety of different assays designed to identify toxic conditions, screen agents for potential toxicity, investigate cellular responses to toxic insults and assay for cell death. A variety of different parameters can be monitored to assess toxicity. Examples of such parameters include, but are not limited to, cell proliferation, monitoring activation of cellular pathways for toxicological responses by gene or protein expression analysis, DNA fragmentation; changes in the composition of cellular membranes, membrane permeability, activation of components of death-receptors or downstream signaling pathways (e.g., caspases), generic stress responses, NF-kappaB activation and responses to mitogens. Related assays are used to assay for apoptosis (a programmed process of cell death) and necrosis.

By contacting various microbial cells with different test compounds, one can also utilize the devices provided herein to conduct antimicrobial assays, thereby identifying potential antibacterial compounds. The term “microbe” as used herein refers to any microscopic and/or unicellular fungus, any bacteria or any protozoan. Some antimicrobial assays involve retaining a cell in a cell cage and contacting it with at least one potential antimicrobial compound. The effect of the compound can be detected as any detectable change in the health and/or metabolism of the cell. Examples of such changes, include but are not limited to, alteration in growth, cell proliferation, cell differentiation, gene expression, cell division and the like.

Certain of the microfluidic devices provided herein can be utilized to conduct mini-sequencing reactions or primer extension reactions to identify the nucleotide present at a polymorphic site in a target nucleic acid. In general, in these methods a primer complementary to a segment of a target nucleic acid is extended if the reaction is conducted in the presence of a nucleotide that is complementary to the nucleotide at the polymorphic site. Often such methods are single base pair extension (SBPE) reactions. Such method typically involve hybridizing a primer to a complementary target nucleic acid such that the 3′ end of the primer is immediately adjacent the polymorphic site, or is a few bases upstream of the polymorphic site. The extension reaction is conducted in the presence of one or more labeled non-extendible nucleotides (e.g., dideoxynucleotides) and a polymerase. Incorporation of a non-extendible nucleotide onto the 3′ end of the primer prevents further extension of the primer by the polymerase once the non-extendible nucleotide is incorporated onto the 3′ end of the primer.

Related to the methods just described, the present devices can also be utilized to amplify and subsequently identify target nucleic acids in multiple samples using amplification techniques that are well established in the art. In general such methods involve contacting a sample potentially containing a target nucleic acid with forward and reverse primers that specifically hybridize to the target nucleic acid. The reaction includes all four dNTPs and polymerase to extend the primer sequences.

While the present invention has been described herein with reference to particular embodiments thereof, a latitude of modification, various changes and substitutions are intended in the foregoing disclosure, and it will be appreciated that in some instances some features of the invention will be employed without a corresponding use of other features without departing from the scope of the invention as set forth. Therefore, many modifications may be made to adapt a particular situation or material to the teachings of the invention without departing from the essential scope and spirit of the present invention. It is intended that the invention not be limited to the particular embodiment disclosed as the best mode contemplated for carrying out this invention, but that the invention will include all embodiments and equivalents falling within the scope of the claims. 

What is claimed is:
 1. A system for analyzing a property of a target material, the system comprising: a structure comprising: a first microfabricated region configured to contain a volume of a first static fluid including a target; a second microfabricated region configured to contain a volume of a second static fluid including an analyte known to bind to the target; a valve actuable to place the volume of the first fluid in diffusive communication with the volume of the second fluid across a free microfluidic interface; and a detector configured to detect a presence of the analyte in the first microfabricated region.
 2. The system of claim 1 wherein: the first microfabricated region comprises a first chamber in fluid communication with a first channel portion restricted in at least one dimension relative to dimensions of the first chamber; the second microfabricated region comprises a second chamber in fluid communication with a second channel portion restricted in at least one dimension relative to dimensions of the second chamber; the valve comprises an elastomer membrane deflectable into the channel between the first channel portion and the second channel portion; and the detector is configured to detect a rate of increase of the concentration of the analyte in the first chamber over time.
 3. The system of claim 1 wherein: the first microfabricated region comprises a first microfabricated channel portion restricted in at least one dimension; the second microfabricated region comprises a second microfabricated channel portion restricted in the at least one dimension; the valve comprises an elastomer membrane deflectable into the microfabricated channel between the first channel portion and the second channel portion; and the detector is configured to detect a concentration of the analyte over a distance from the valve in the first chamber.
 4. The system of claim 1 wherein the detector is configured to detect at least one of fluorescence, refractive index, colorimetry, surface plasmon resonance, and conductivity. 